Molecular wire injection sensors

ABSTRACT

Disclosed is a sensor for sensing the presence of an analyte component without relying on redox mediators. This sensor includes (a) a plurality of conductive polymer strands each having at least a first end and a second end and each aligned in a substantially common orientation; (b) a plurality of molecular recognition headgroups having an affinity for the analyte component and being attached to the first ends of the conductive polymer strands; and (c) an electrode substrate attached to the conductive polymer strands at the second ends. The electrode substrate is capable of reporting to an electronic circuit reception of mobile charge carriers (electrons or holes) from the conductive polymer strands. The electrode substrate may be a photovoltaic diode.

CROSS-REFERENCE TO RELATED APPLICATIONS

[0001] This application is a continuation, claiming priority under 35U.S.C. §120, from U.S. patent application Ser. No. 09/960,165, filedSep. 20, 2001 naming inventor Randy E. Keen, and titled “MOLECULAR WIREINJECTION SENSORS,” which is a continuation-in-part U.S. applicationSer. No. 08/856,822, filed May 14, 1997, now U.S. Pat. No. 6,060,327,issued May 9, 2000. Both applications are incorporated herein byreference for all purposes.

BACKGROUND OF THE INVENTION

[0002] The present invention relates to biosensors and chemical sensors.More particularly, it relates to sensors having a chemical orbiochemical species detection group connected to an electronic circuitby electrically conducting polymer strands.

[0003] Biosensors employing enzymes have been applied to the detectionof numerous analyte species concentrations including glucose,cholesterol, or both glucose and cholesterol concentrations in wholeblood samples. Such sensors and associated instruments employ an enzymecapable of catalyzing a reaction at a rate representative of theselected compound concentration in an assay mixture.

[0004] There are three general detection approaches employing a glucoseenzyme electrode. The first and earliest measures oxygen consumption.The oxygen-sensing probe is an electrolytic cell with a gold (orplatinum) cathode separated from a tubular silver anode by an epoxycasting. The anode is electrically connected to the cathode byelectrolytic gel, and the entire chemical system is isolated from theenvironment by a thin gas-permeable membrane (often Teflon). A potentialof approximately 0.8 V (from solid-state power supply) is appliedbetween the electrodes. The oxygen in the sample diffuses through themembrane and is reduced at the cathode with the formation of theoxidation product, silver oxide, at the silver anode. The resultantcurrent is proportional to the amount of oxygen reduced. The analyzerunit operates over the range from 0.2 to 50 ppm of dissolved oxygen.Gases that reduce at 0.8 V will interfere; these include the halogensand SO₂. H₂S contaminates the electrodes.

[0005] A second approach detects H₂O₂ production but requires an appliedpotential of approximately 0.65 V (from solid-state power supply)applied between the electrodes, one of which is inside a permselectivemembrane. The H₂O₂ in the sample diffuses through the permselectivemembrane (if one is present) and is oxidized at the anode. Many metal,metal complexes, nonmetal, organic and biochemical species that oxidizeat approximately 0.65 V will interfere; such as ascorbic acid, amines,hydrazines, thiol compounds, catechols, hydroquinones, ferrocenes, andmetalloporphyrins. The inside permselective membrane is not alwayscapable of removing the complicated mix of possible interferences fromthe analyte matrix.

[0006] A third approach takes advantage of the fact that the enzymaticreaction requires two steps. First, the enzyme glucose oxidase (GOD) (EC1.1.3.4) is reduced by glucose, then the reduced enzyme is oxidized toits initial form by an electron acceptor, i.e., a mediator. In naturalsystems, the mediator is oxygen. In biosensors, another mediatorcompound may be employed to transfer electrons between the enzyme and aconductive surface of an electrode at a rate representative of theenzyme catalyzed reaction rate when an appropriate potential is appliedto the particular redox mediator in use. Such biosensors may employamperometric measurements to determine glucose concentration in a wholeblood sample. This involves an integrated sample measurement of the areaunder the ampere versus time curve, corresponding to the amount ofglucose in the sample.

[0007] The mechanism by which a common amperometric sensor works isdepicted in FIG. 1. A sensor 2 employs glucose oxidase (GOD), forexample, as a molecular recognition group. Glucose oxidase catalyzes theoxidation of glucose to gluconolactone in analyte 4. This reactioninvolves the FAD/FADH₂ redox center of the enzyme. Sensor 2 includes amolecular recognition group, region 6, attached to an electrode 8. Whenglucose in analyte 4 contacts GOD-FAD (glucose oxidase including the FADredox center) in region 6, it is oxidized to gluconolactone. At the sametime, the GOD-FAD is reduced to GOD-FADH₂. This involves two electronsand two hydrogen ions being transferred to the FAD. Normally, in theabsence of a sensor mediator, the GOD-FADH₂ is reoxidized by atmosphericoxygen to GOD-FAD to complete the catalytic reaction. In the presence ofa mediator, however, the GOD-FADH₂ is sometimes reoxidized by a mediator(Mox). In this case, the GOD-FADH₂ releases two hydrogen ions to analyte4 and two electrons to the mediator. The resulting reduced mediator(Mred) may then be reoxidized by electrode 8 at an appropriatepotential. The reoxidation of the mediator is accompanied by thetransfer of an electron or electrons to electrode 8. This is the currentthat is monitored to provide a concentration of glucose.

[0008] In theory, a mediator may be any small molecule inorganic,organometallic or organic compounds, which are reduced by the enzyme,and oxidized by an appropriate applied potential at the electrodesurface. The mediator should be designed to rapidly and efficientlytransfer electrons between the enzyme and the electrode. Otherwise,ambient oxygen would oxidize nearly all of the reduced GOD and thedesired signal would be very weak. The mediator should also transfer atotal charge proportional to the glucose or cholesterol concentration inthe sample. The current which results from the mediator oxidation isknown as the Cottrell current which, when integrated with respect totime, gives the number of coulombs associated with the sensor reaction.The total coulombs passed is proportional to the amount of analyte.

[0009] Unfortunately, mediators are commonly provided as mobile“reagents” which diffuse to the enzyme where they are oxidized orreduced (depending upon the reaction catalyzed by the enzyme). Theoxidized or reduced mediator then diffuses to the electrode surfacewhere it gains or loses an electron. Unfortunately, such mechanism isdependent upon the continuing presence of recycled mobile mediators. Assuch compounds can leak from the electrode surfaces, there may be agradual depletion in available mediator and a consequent reduction insensor sensitivity. Examples of diffusing redox mediators include dyes(e.g., methylene blue), ferrocene derivatives (Cass, A E G; Davis, G;Francis, G D; Hill, H A O; Aston, W J; Higgins, I J; Plotkin, E V;Scott, L D L; Turner, A P F: Ferrocene-Mediated Enzyme Electrode forAmperometric Determination of Glucose. Anal. Chem. 56:667-671, 1984),components of conducting organic metals and quinones.

[0010] Also, available sensors applying the above amperometric approachto the detection of glucose, cholesterol, lactate, H₂O₂, NAD(P)H,alcohol, and a variety of other compounds in whole blood samples, canhave other serious complicating problems. For example, the percentage ofsensor surface area covered by blood can vary; sometimes the bloodsample does not cover the entire electrode. This may be caused by apoorly adherent enzyme (often applied by spraying) thus allowing leakageof blood or other analytes along the edges of the electrode. A relatedproblem results from hydration of the reaction area prior to test. Thisdilutes the ligand (e.g., glucose) concentration and therefore gives alower reading than would be accurately given by an unhydrated surface.

[0011] Further, the partial pressure of molecular oxygen (O₂) maycomplicate the interpretation of sensor data. Molecular oxygen is thenatural electron acceptor mediator of the enzyme glucose oxidase (GOD).Following oxidation of D-(+)-glucose by GOD, reduced glucose oxidase(GODred) will transfer electrons to O₂ forming H₂O₂ in the absence ofother mediators. In amperometric glucose biosensors described above, theunwanted O₂ side reaction competes with synthetic chemical mediators forelectrons supplied by the GODred enzyme. Calibration of GOD-basedbiosensors at different altitudes (i.e., different partial pressures ofO₂) may be a problem if electron transfer rates of selected syntheticchemical mediators are not orders of magnitude faster than the O₂ sidereaction.

[0012] Humidity (i.e., H₂O) may be another potential problem if massaction of H₂O and O₂ present drives the enzyme catalyzed oxidationproduct of D-gluconolactone in reverse back to the reduced startingmaterial, D-(+)-glucose. Catalase, a common contaminant of glucoseoxidase preparations, may be driven in reverse by mass action of excessH₂O and O₂ producing 2 moles of H₂O₂. H₂O₂ buildup combined withD-gluconolactone could drive the glucose oxidase reaction in reverse bymass action back to D-(+)-glucose.

[0013] Other problems associated with known amperometric sensorsinclude, for example, (1) difficulty in fitting the Cottrell currentcurve (i.e., ampere-time graph), (2) sampling with enough frequency toaccurately obtain the time integral of Cottrell current, (3) highapplied potential at the electrode causing indiscriminate oxidation orreduction of interfering substances, and (4) complicated electroniccircuits requiring potentiostat and galvinostat instrumentation.

[0014] Some of the above drawbacks of the current amperometricbiosensors have been noted and analyzed (see, Schuhmann, W: Chap. 9.Conducting Polymers And Their Application In Amperometric Biosensors.In: Diagnostic Biosensor Polymers. ACS Symposium Series 556. Usmani, AM; Akmal, N; eds. American Chemical Society; Washington, D.C.; 1994; pp.110-123). First, due to the fact that the active site of redox enzymesis in general deeply buried within the protein shell, direct electrontransfer between enzymes and electrode surfaces is rarely encountered.This is especially true for enzymes which are integrated withinnon-conducting polymer membranes in front of the electrode surface.Hence, electron transfer is usually performed according to a ‘shuttle’mechanism involving free-diffusing electron-transferring redox speciesfor example the natural electron acceptor O₂ or artificial redoxmediators like ferrocene derivatives (Cass, A E G; Davis, G; Francis, GD; Hill, H A O; Aston, W J; Higgins, U; Plotkin, E V; Scott, L D L;Turner, A P F: Ferrocene-Mediated Enzyme Electrode for AmperometricDetermination of Glucose. Anal. Chem. 56:667-671, 1984), osmiumcomplexes (Heller, A: Electrical Wiring of Redox Enzymes. Acc. Chem.Res. 23(5):128-134, 1990), or quinones. Due to the necessity for theredox mediators to diffuse freely between the active sites of theenzymes and the electrode surface, these electrodes show a limitedlong-term stability as a consequence of the unavoidable leaking of themediator from the sensor surface. Additionally in the case of thenatural redox couple O₂/H₂O₂, the sensor signal is dependent on the O₂partial pressure, and a high operation potential has to be applied tothe working electrode giving rise to possible interferences fromcooxidizable compounds. The second drawback is related to thefabrication of these sensors. The physical assembling of an enzymemembrane and an electrode is extremely difficult to automate and thus inprincipal incompatible with microelectronic fabrication techniques.Additionally, the miniaturization as well as the integration ofindividual biosensors into a miniaturized sensor array is impossiblewith techniques which are mainly based on the manual deposition of adroplet of the membrane-forming mixture onto the electrode surface.

[0015] Consequently, the next generation of amperometric enzymeelectrodes has to be based on immobilization techniques which arecompatible with microelectronic mass-production processes and easy tominiaturize. Additionally, the integration of all necessary sensorcomponents on the surface of the electrode has to prevent the leaking ofenzymes and mediators simultaneously improving the electron-transferpathway from the active site of the enzyme to the electrode surface.

[0016] In addition to amperometric mechanisms, which rely on detectingcurrent generated from faradaic reactions, a potentiometric mechanismmay be employed to sense analyte concentration. Potentiometrictechniques monitor potential changes between a working electrode and areference electrode in response to charged ion species generated fromenzyme reactions on the working electrode. A very common potentiometricsensor is the pH sensor which registers changes in hydrogen ionconcentration in an analyte. A microelectronic potentiometric biosensor,the Field Effect Transistor (FET) biosensor, has generated someinterest. In this design, a receptor or molecular recognition species iscoated on a transistor gate. When a ligand binds with the receptor, thegate electrode potential shifts, thereby controlling the current flowingthrough the FET. This current is detected by a circuit which converts itto an observed ligand concentration. Observed problems withpotentiometric systems include, for example, (1) slow response of theelectrode (i.e., seconds), (2) complicated electronic circuits for threeelectrode (i.e., working, counter, and reference electrode)electrochemical systems requiring potentiostat instrumentation, (3) lowsensitivity, and (4) limited dynamic range.

[0017] Recently, two groups (Heller et al. and Skotheim et al.) haveexplored and developed redox polymers that can shuttle electrons fromthe enzyme to the electrode. The groups have “wired” the enzyme to theelectrode with a long redox polymer having a dense array of electronrelays. Each relay is a redox site bound to the polymer backbone.Electrons move along the polymer by hopping from one redox appendage tothe next. The polymer penetrates and binds the enzymes, and is alsobound to the electrode.

[0018] Heller et al. have conducted work on Os-containing redoxpolymers. They have synthesized a large number of such Os-containingpolymers and evaluated their electrochemical characteristics (Gregg, BA; Heller, A: Redox Polymer Films Containing Enzymes. 1. ARedox-Conducting Epoxy Cement: Synthesis, Characterization, andElectrocatalytic Oxidation of Hydroquinone. J. Phys. Chem. 95:5970-5975,1991). Their most stable and reproducible redox polymer is apoly(4-vinyl pyridine) to which Os(bpy)₂Cl₂ has been attached to ⅙th ofthe pendant pyridine groups. The resultant redox polymer is waterinsoluble. To make it water soluble and biologically compatible, Helleret al. have partially quaternized the remaining pyridine pendants with2-bromoethyl amine. The redox polymer is water soluble and the newlyintroduced amine groups can react with a water soluble epoxy e.g.,polyethylene glycol diglycidyl ether and GOD to produce a cross-linkedbiosensor coating-film. Such coating-films produced high currentdensities and a linear response to glucose up to 600 mg/dL (U.S. Pat.No. 5,262,035 to Gregg et al.).

[0019] Heller describes the electrical wiring of redox enzymes for useas amperometric biosensors (Heller, A: Electrical Wiring of RedoxEnzymes. Acc. Chem. Res. 23(5):128-134, 1990). The Heller approach is animprovement over amperometric enzyme electrodes based on diffusing redoxmediators, including dyes, ferrocene derivatives, components ofconducting organic metals, and quinones, all described above. In theHeller approach, redox centers of a redox polymer polycation (e.g.,2[Os-(2,2′-bipyridine)₂(poly(vinylpyridine))Cl]^(1+/2+)) areelectrostatically and covalently bound to the enzyme and relayselectrons to the electrode, on which a segment of the polycation isadsorbed. Binding of the redox polymer polycation to the electrode canbe electrostatic when the electrode has a negative surface charge.

[0020] Fluctuations in current with partial pressure of oxygen (e.g.,oxygen concentration in blood), depend on the ratio of the rate ofdirect electroxidation of the FADH₂ centers to their rate of oxidationby molecular oxygen, and therefore on the rate of electron transfer to,and the electrical resistance of, the three-dimensional wired-enzymestructure. At high osmium-complex concentrations, and in sufficientlythin layers, the competition is won by electron transfer to theelectrode via the osmium centers, and the electrodes are relativelyinsensitive to oxygen (Heller, A: Electrical Wiring of Redox Enzymes.Acc. Chem. Res. 23(5):128-134, 1990. Gregg, B A; Heller, A: Cross-LinkedRedox Gels Containing Glucose Oxidase for Amperometric BiosensorApplications. Anal. Chem. 62:258-263, 1990. Surridge, N A; Diebold, E R;Chang, J; Neudeck, G W: Chap 5. Electron-Transport Rates In An EnzymeElectrode For Glucose. In: Diagnostic Biosensor Polymers. ACS SymposiumSeries 556. Usmani, A M; Akmal, N; eds. American Chemical Society;Washington, D.C.; 1994; pp. 47-70).

[0021] Electrodes based on conducting polypyrroles with ferrocenes alsohave been reported (Hale, P D; Inagaki, T; Karan, H I; Okamoto, Y;Skotheim, T A: A New Class of Amperometric Biosensor Incorporating aPolymeric Electron-Transfer Mediator. J. Am. Chem. Soc.111(9):3482-3484, 1989).

[0022] Skotheim et al. have used flexible polymer chains to act asrelays. Their polymers provide communication between GOD's redox centersand electrode. No mediation was found when ferrocene was attached to anon-silicone backbone. Their ferrocene-modified siloxane polymers weresaid to be stable and non-diffusing (Boguslavsky, L I; Hale, P D;Skotheim, T A; Karan, H I; Lee, H S; Okamoto, Y: Novel Biosensors ForSpecific Neurotransmitters Based On Flavoenzymes And Flexible RedoxPolymers. Polym. Mater. Sci. Eng. 64:322-323, 1991).

[0023] Unfortunately, the redox polymer systems of Heller et al. andSkotheim et al. have a limited electron transfer rate based on electronhopping between dense electron relay pendant groups. Further, their“wire” redox centers must be designed to undergo reaction at a potentialclose to that of the enzyme catalyzed reaction. The closer the potentialis to the redox potential of the enzyme itself, the lesser thelikelihood that a potentially interfering substrate will be spuriouslyoxidized. Unfortunately, to address this issue limits the range ofpolymer redox couple and molecular headgroup combinations.

[0024] A fundamental presupposition for the construction of reagentlessamperometric enzyme electrodes is the design of a suitableelectron-transfer pathway from the active site of the enzyme to theelectrode surface. According to Marcus theory (Marcus, R A; Sutin, N:Electron Transfers In Chemistry And Biology. Biochim. Biophys. Acta811:265-322, 1985) a redox mediator with a low reorganization energyafter the electron transfer has to be able to penetrate into the activesite of the enzyme to shorten the distance between the prosthetic group(e.g., FAD/FADH₂) and the mediator. Hence, the rate constant of theelectron-transfer reaction can be increased. After this ‘first’ electrontransfer the redox equivalents have to be transported to the electrodesurface via mechanism having a rate constant which is in the range ofthe turnover rate of the enzyme. In the shuttle mechanism mentionedabove (having mobile mediators), the electron transport involvesdiffusion of redox mediators. In the “wired” redox polymer sensorsdescribed above, electron transport involves hopping from one redoxcenter to the next on the polymer backbone.

[0025] In a recent study, Aizawa et al. discuss a reversible electrontransfer between the prosthetic group of pyrrolo quinoline quinone (PQQ)enzyme (fructose dehydrogenase) and an electrode through a molecularinterface (Aizawa, M; Khan, G F; Kobatake, E; Haruyama, T; Ikariyama, Y:Chap. 26. Molecular Interfacing of Enzymes on the Electrode Surface. In:Interfacial Design and Chemical Sensing. ACS Symposium Series 561.Mallouk, T E; Harrison, D J; eds. American Chemical Society, Washington,D.C., 1994, pp. 305-313). The PQQ moieties of randomly oriented fructosedehydrogenase (FDH) which are very close to the transducer electrode caneasily transfer their electrons to the electrode (Shinohara, H; Khan, GF; Ikariyama, Y; Aizawa, M: Electrochemical Oxidation and Reduction ofPQQ Using a Conducting Polypyrrole-Coated Electrode. J. Electroanal.Chem. 304:75-84, 1991. Khan, G F; Shinohara, H; Ikariyama, y; Aizawa, M:Electrochemical Behaviour of Monolayer Quinoprotein Adsorbed on theElectrode Surface. J. Electroanal Chem. 315:263-273, 1991). However, theprosthetic groups of FDH located far from the electrode can not providetheir electrons, as the distance from the electrode exceeds the maximumelectron transfer distance (˜25 Å). Therefore, to make the FDH (EC1.1.99.11, MW: 141,000) on the electrode surface electrochemicallyactive, Aizawa et al. introduced an ultrathin conductive polypyrrole(PP) membrane as a molecular interface as “wiring” to assist theelectron transfer from PQQ to the electrode. Unfortunately, the wiringused by Aizawa is randomly oriented and does not necessarily presentenzyme at optimal position with respect to the analyte.

[0026] What is needed is an improved sensor design that rapidlytransfers electrons from headgroup redox reactions to an electrode, doesnot rely on a redox relay such as freely diffusing mediators, andoptimally orients the headgroup with respect to the analyte.

[0027] A great number of approaches for microfabrication of chemicalsensors are currently under way, particularly in the areas of fieldeffect transistor (FET)-based chemical sensors, metal oxide gas sensors,and biosensors. Since Janata et al. first reported micro-enzymeelectrodes based on FET (Caras, S; Janata, J: Field Effect TransistorSensitive to Penicillin. Anal. Chem. 52:1935-1937, 1980), a number ofgroups have been employing microfabrication techniques (e.g.,photolithography) such as those employed in semiconductor devicetechnology to fabricate micro-enzyme electrodes. Despite enormousefforts of many groups, the FET-based micro-enzyme electrodes ofpractical use have not been realized yet, largely because of theproblems associated with potentiometric methods general lack of a fastresponse, high sensitivity, and wide dynamic range.

[0028] For the construction of reagentless enzyme electrodes (e.g.,electrodes analogous to those of Heller et al. and Aizawa et al.) onehas to focus on a technique for the modification and functionalizationof electrode and even micro-electrode surfaces to allow the strongbinding of the enzyme and the redox mediator taking into account thepresuppositions for an effective and fast electron transfer between theenzyme and the electrode. These features requirements are in principlemet with enzyme electrodes based on redox-sensitive hydrogels, however,the manual deposition of these hydrogels is not compatible withmass-production techniques.

[0029] The electrochemical deposition of conducting-polymer layersoccurs exclusively on the electrode surface and can hence be used forthe immobilization of enzymes either covalently using functionalities onthe polymer film or physically entrapped within the growing polymerfilm. As the conducting-polymer film itself does not participate in theelectron transfer, mediator-modified enzymes entrapped within apolypyrrole layer have been used for the construction of a reagentlessoxidase electrode.

[0030] Electrochemical deposition methods of the prior art typically usehigh current density and voltage potential conditions which destroy theorderly Helmholtz double-layer at the electrode surface (U.S. Pat. No.5,215,631 to Westfall). Resulting disorderly depositions at electrodesurfaces produce random polymer structures which lack orientational andpositional order. Aizawa et al. “wired” PQQ-FDH in their sensors withultrathin conductive polypyrrole (PP) membrane as a molecular interface.Electrochemical synthesis of molecular-interfaced FDH on Pt electrodewas prepared by the following two steps: (1) potential-controlledadsorption of FDH, and (2) electrochemical polymerization ofpolypyrrole. These steps employ high voltage and current densityelectrochemical deposition conditions to produce polymer (FDH andpolypyrrole) depositions on the Pt electrode that are randomly oriented.Therefore, this device must operate at high (˜400 mV) operatingpotential resulting in possible interfering cooxidizable species.

[0031] What is needed is an improved technique for depositing molecularrecognition groups and associated wiring, if necessary, that provides astrong direct connection between an electrode and the molecularrecognition groups, and allows the molecular recognition groups to bealigned in a common orientation.

SUMMARY OF THE INVENTION

[0032] In one aspect, the present invention provides a sensor forsensing the presence of an analyte component without relying on redoxmediators. This sensor may be characterized as including the followingelements: (a) a plurality of conductive polymer strands each having atleast a first end and a second end and each aligned in a substantiallycommon orientation; (b) a plurality of molecular recognition headgroupshaving an affinity for the analyte component and being attached to thefirst ends of the conductive polymer strands; and (c) an electrodesubstrate attached to the conductive polymer strands at the second ends.

[0033] The polymer strands in a common orientation resemble liquidcrystals. Preferably, the strands are oriented substantially orthogonalto the electrode substrate. The conductive polymer strands may be, forexample, one or more of multi-stranded nucleic acids, electron transportproteins, synthetic organic and inorganic conducting polymers, metalcrystallite molecular wires, and Langmuir-Blodgett conducting films. Ina particularly preferred embodiment, the conductive polymer strands aredouble-stranded DNA strands.

[0034] The headgroup may participate in a redox reaction when contactinga molecule of the analyte component. When this is the case, a mobilecharge carrier is transferred directly to a conductive polymer strandattached to the headgroup, without participating in a redox reaction inthe polymer strand. In one embodiment, the molecular recognitionheadgroups participate in the redox reaction by catalyzing a chemicaltransformation of the analyte component. Examples of such headgroupsinclude oxidoreductases and catalytic antibodies. In one specificexample used repeatedly in this specification, the headgroup is glucoseoxidase.

[0035] The sensor headgroups may be chemically homogeneous (e.g., theyare all glucose oxidase) or chemically inhomogeneous (e.g., they includea mixture of glucose oxidase, cholesterol oxidase, and cholesterolesterase). In one preferred embodiment, when the headgroups areinhomogeneous, the sensor includes a first region on the electrodesubstrate where a first group of chemically homogeneous molecularrecognition headgroups is located and second region on the electrodesubstrate where a second group of chemically homogeneous molecularrecognition headgroups is located. The first and second regions may beseparately addressable so that information signal from the two regionsmay be separately processed and able to indicate whether cholesterol,glucose, or both cholesterol and glucose are present in the analyte forexample.

[0036] The electrode substrate should be capable of reporting to anelectronic circuit reception of mobile charge carriers from theconductive polymer strands. In one specific embodiment, the electrodesubstrate is a diode such as a photovoltaic diode. More generally, thesubstrate may be a device element of a device on semiconductor chip(e.g., a gate on an FET).

[0037] In a variation of this aspect of the invention, a sensor isprovided to detect the presence of a nucleic acid sequence (at a crimescene for example). The sensor includes (a) a plurality ofsequence-specific single-stranded nonconductive nucleic acid wires eachhaving at least a first end and a second end; and (b) an electrodesubstrate attached to sequence-specific single-stranded nonconductivenucleic acid strands at the second ends and capable of reporting to anelectronic circuit, reception of mobile charge carriers originating fromcomplementary multi-stranded nucleic acid strands. In this embodiment,when the sensor is exposed to an analyte having the complementarynucleic acid sequence, at least some of the affixed single-strandednonconductive nucleic acid wires hybridize or anneal with the analyte toform conductive multi-stranded nucleic acid strands. Thus, chargecarriers can be transported to the electrode substrate for detection. Inone embodiment, the plurality of sequence-specific single-strandednonconductive nucleic acid strands are attached to molecular recognitionheadgroups such that mobile charge carriers are transferred directlythrough only annealed multi-stranded nucleic acid strands when a redoxreaction occurs at the attached molecular recognition headgroups.

[0038] Another aspect of the invention provides method of detecting aconcentration of an analyte component in an analyte with a sensor havinga structure as described above. The method may be characterized asincluding the following steps: (a) contacting the molecular recognitionheadgroups with the analyte; and (b) determining whether electrons havebeen transferred to the electrode substrate resulting from electronsgenerated by the redox reaction and transferred by the conductivepolymer strands to the electrode substrate. When the redox reactionoccurs at a headgroup, a mobile charge carrier is transferred directlyto a conductive polymer strand attached to the headgroup, without redoxreaction in the polymer strand. The method may further involve (c)monitoring a change in an electronic circuit connected to the electrodesubstrate, the change resulting from reception of mobile charge carriersfrom the conductive polymer strands; and (d) correlating the change inthe electronic circuit with the concentration of the analyte component.

[0039] Another important aspect of the claimed invention is a sensoremploying a diode, preferably a photodiode. Sensors in accordance withthis aspect of the invention may be characterized as including thefollowing features: (a) a plurality of molecular recognition headgroupshaving an affinity for the analyte component and participating in aredox reaction when contacting a molecule of the analyte component suchthat when the redox reaction occurs at a headgroup, a mobile chargecarrier is generated; (b) a diode having a first electrode to which theplurality of molecular recognition headgroups are affixed such thatmobile charge carriers generated by the redox reaction are transferredto the first electrode; and (c) a circuit for detecting when the mobilecharge carriers are transferred to the first electrode. In a preferredembodiment, the plurality of molecular recognition headgroups areattached to a p-type side of the diode. Also the diode may be a deviceon semiconductor chip including a plurality of devices.

[0040] In a further preferred embodiment, the headgroups are attachedthrough conductive polymer strands arranged as described in the aboveembodiments. Thus, for example, the conductive polymer strands may besubstantially commonly oriented (e.g., orthogonal to the diode surface).

[0041] A diode sensor as described above may be used according to amethod as follows: (a) contacting the molecular recognition headgroupswith the analyte; (b) specifying a baseline electrical signal that ispresent when (i) a stimulus is provided to the diode and (ii) theplurality of molecular recognition headgroups are substantially free ofthe analyte component; and (c) detecting a deviation from the baselineelectrical signal, which deviation results from transfer of the mobilecharge carriers to the first electrode when the analyte component comesin contact with the molecular recognition headgroups. The method mayfurther include (d) determining an amplitude of the deviation; and (e)determining an analyte component concentration directly from theamplitude of the deviation without the use of any other information fromthe electrical signal. It has been found that the analyte componentconcentration is sometimes proportional to the amplitude of thisdeviation. Depending upon the type of signal detector employed, thebaseline electrical signal and the deviation from the baselineelectrical signal may be measures of voltage or electrical current.Preferably, though not necessarily, the diode is a photovoltaic diodeand the stimulus provided in the specifying a baseline electrical signalis radiant energy.

[0042] Yet another aspect of the present invention is method of forminga sensor capable of sensing the presence of an analyte component. Thismethod may be characterized as including the following: (a) contacting asensor substrate (e.g., a device element of a device on semiconductorchip) with a first medium containing mobile conductive polymer strandsor precursors of the conductive polymer strands; (b) applying a firstpotential to the substrate sufficient to form a first structure havingthe conductive polymer strands affixed to the substrate; (c) contactingthe sensor substrate, with affixed conductive polymer strands, with asecond medium containing mobile molecular recognition headgroups; and(d) applying a second potential to the substrate sufficient to affix themolecular recognition headgroups to the affixed conductive polymerstrands. This process produces a sensor structure in which the substrateaffixed to the conductive polymer strands and the molecular recognitionheadgroups also affixed to the conductive polymer strands.

[0043] Preferably, the step of applying a first potential is performedat a potential which causes the affixed conductive polymer strands to beoriented in a substantially common direction. This potential may bebetween about 0.001 and 500 mV, for example. The step of applying asecond potential is preferably performed at a potential which causes theaffixed molecular recognition headgroups to be oriented in asubstantially common direction. This second potential may be betweenabout 0.001 and 500 mV. Preferably, though not necessarily, the firstmedium is removed from the sensor substrate following the step ofapplying a first potential. In an alternative embodiment, the secondmedium is obtained from the first medium by performing the step ofapplying a first potential.

[0044] If a sensor having separated regions of different headgroups isto be created, the method may also require isolating a region of thesensor substrate prior to the step of contacting the sensor substratewith a second medium, such that the molecular recognition headgroups aredeposited only in the isolated region. To produce multiple headgroupregions, the steps of isolating a region, contacting the sensorsubstrate with a second medium, and applying a second potential to thesubstrate are performed a second time. The step of contacting the sensorsubstrate with a second medium for a second time employs a secondmolecular recognition headgroup, to form a structure having a firstregion on the sensor substrate having a first group of chemicallyhomogeneous molecular recognition headgroups and a second region on thesensor substrate having a second group of chemically homogeneousmolecular recognition headgroups.

[0045] Sensors of this invention provide analyte concentration readings,fast responses, high sensitivity, high dynamic range, and few erroneousreadings. In a glucose sensor of this invention, glucose concentrationis accurately read despite changes in partial pressure Of O₂,atmosphere, altitude, humidity, or sample application of blood.Specifically, the direct wired enzyme sensors of the present inventionovercome the difficulty caused by molecular oxygen reoxidizing a reducedenzyme before that enzyme (or more precisely its redox center) canrelease electrons to the electrode. This is because the directly wiredsensors of this invention may provide electron transfer rates manyorders of magnitude faster than enzymatic reaction rates, and electrontransfer rates of diffusional redox mediators such as O₂ and otherartificial mediators. This provides sub-millisecond digital output fromthe sensing chip.

[0046] Chips based on device molecular transistors may be reusable,disposable, reagentless, membraneless. Further, they are amenable tominiaturization and mass production, do not require complicated threeelectrode systems (i.e., no working, counter, or reference electrodes)and associated electrochemical instrumentation (i.e., no galvinostat orpotentiostat), and provide real-time digital output directly from thechip. These and other features and advantages of the present inventionwill be described in more detail below with reference to the drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

[0047]FIG. 1 is a representation of the mechanisms employed in aconventional redox mediator based biosensor.

[0048]FIG. 2 is a representation of a sensor-solution interface inaccordance with this invention and showing a substrate, molecular wire,and molecular recognition headgroup.

[0049]FIG. 3 is a schematic illustration of photodiode sensor inaccordance with an embodiment of the present invention.

[0050]FIG. 4A is representation of an electrodeposition step forattaching molecular wires to a substrate in accordance with anembodiment of this invention.

[0051]FIG. 4B is representation of an electrodeposition step forattaching molecular recognition headgroups to molecular wires (depositedas shown in FIG. 4A) in accordance with an embodiment of this invention.

[0052]FIG. 5 is a graph showing a current signal generated when glucoseis contacted with a photodiode type GOD glucose sensor in accordancewith one embodiment of this invention.

[0053]FIG. 6 is a graph showing current and voltage signals generatedwhen the sensor employed in FIG. 5 is subjected to a regimen includingcontact with glucose, washing, open circuit, and recontact with glucose.

[0054]FIG. 7 is a graph showing current and voltage signals generatedfrom a sensor employing GDH on a photodiode when exposed to glucose.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0055] I. Overview

[0056] II. Solid Substrate

[0057] III. Sequential Electrochemical and Chemical DepositionTechniques

[0058] A. Electrochemical Atomic Layer Epitaxy (ECALE)

[0059] B. Sequential Monolayer Electrodeposition (SMED)

[0060] C. Thin Film Chemical Deposition (CD)

[0061] D. Electrochemical Molecular Layer Epitaxy (EMOLE)

[0062] 1. Deposition of Uniaxially Oriented Liquid Crystal ConductingBiopolymers (Proteins and DNA)

[0063] IV. Conducting Polymers and Thin Films

[0064] A. Electron Transport Proteins

[0065] B. DNA Quantum Wires

[0066] V. Molecular Recognition Surfaces

[0067] A. Oxidoreductases (Redox Enzymes)

[0068] B. Immunoglobulins

[0069] VI. Conduction Mechanisms through Polymers on Solid Substrates

[0070] A. Energy Bands in Uniaxially Oriented Liquid Crystal ConductingBiopolymers (Proteins and DNA) and Semiconductor Substrates

[0071] B. Superconductivity

[0072] VII. Applications

[0073] VIII. Screening and Assays

[0074] IX. Examples

[0075] I. Overview

[0076] The present invention relates to sensors, sensor fabricationprocesses and semiconductor devices that include the sensors. Thesensors and related devices may be used for recognizing the presence of,quantitating the amount of, and/or continuously monitoring the level of,one or more selected components in a solid, semi-solid, liquid, or gasmixture. Preferably, an active molecular recognition surface is “hardwired” to the substrate surface (e.g., a semiconductor surface) by anoriented liquid crystal wire that is itself conductive. The molecularrecognition surface may be of biologically active material of the typeconventionally employed in sensors. The substrate may be patterned orunpatterned and may include (particularly when semiconductors areinvolved) a conductive coating such as a metal between the underlyingbulk substrate and the liquid crystal wire.

[0077] Hard wiring as that term is used herein may be achieved, in oneembodiment, via electrochemical fabrication methods described in detailbelow. Generally, such methods make use of low-cost, rapid-prototypingsequential electrochemical and chemical deposition techniques such aselectrochemical molecular layer epitaxy (EMOLE) which perform “molecularwiring” and “molecular soldering” procedures. The liquid crystal wiringarrangement preferably provides a “lawn” of commonly oriented “moleculardevices” each including a single molecular recognition site “headgroup”and attached molecular wire “tail.” For context, each such device mightrange in size from about ˜2 to 2500 Å² surface area (e.g., enzyme,enzyme co-factor, substrate, supramolecular assembly, cavitand,host-guest complex, ligand, receptor, antibody, antigen, etc.).

[0078] Biosensors of the present invention may require very lowoperating potentials. In a preferred embodiment, extended conformationof straight uniaxially oriented liquid crystal DNA wires are stuck intothe GOD active site/redox center of the prosthetic group FAD/FADH₂, toprovide an electron transfer pathway to the surface of a p-nhomojunction semiconductor solar cell substrate. A pair of electrons perenzyme turnover event injected from the wires combine with a pair ofholes in the p-type semiconductor layer, interfering with the normalphotocurrent (i.e., electron/hole pair recombination) occurring in thesolar cell. The oriented liquid crystal enzyme (molecular recognitionheadgroup) and attached oriented liquid crystal DNA wire tail constitutea molecular transistor. The device communicates with a solid substrate(i.e., p-n homojunction) through the uniaxially oriented liquid crystalDNA wire tail interconnects. One end of the DNA wire is stuck in theoriented liquid crystal enzyme active site/redox center and the otherend is stuck into the p-type semiconductor layer providing a directconnection between the protein enzyme, DNA, and semiconductor substrate.

[0079] In general, the sensors of this invention may be categorizedbased upon their transduction and/or gating mechanisms of theheadgroup(s): switched or gated by optical (optoelectronic), chemical(chemoelectronic), magnetic (magnetoelectronic), radioactive(radioelectronic), thermal (thermoelectronic), mechanical(piezoelectronic), or electrical (voltage, current, resistivity,capacitance).

[0080]FIGS. 2 and 3 depict sensors structures in accordance with certainpreferred embodiments of the present invention. FIG. 2 presents across-sectional view of a surface region of a sensor 12. As shown,sensor 12 includes an electrode 14 which is preferably made from siliconor another semiconductor substrate. Attached to electrode 14 is aplurality of conducting polymer strands 16. In a preferred embodiment,each strand is a DNA double-stranded molecule. Conductive polymerstrands 16 are orientated substantially in a common direction which isshown to be normal (orthogonal) to substrate 14. Strands 16 are coupledto substrate 14 in a manner that allows direct electrical influencebetween these two features in the sensor. For example, the connectionmight allow electrons to be directly transferred from strands 16 tosubstrate 14 so that circuitry coupled to substrate 14 can detectinjection of electrons. In addition, a potential applied to substrate 14may influence the physical state of conductive polymer strands 16.

[0081] As will be described in more detail below, a preferred processfor affixing polymer strands 16 to substrate 14 provides this directelectronic coupling and in addition orients the strands 16 along asubstantially common axis. Because strands 16 are oriented in asubstantially common direction, they will sometimes be collectivelycharacterized herein as a liquid crystal.

[0082] Note that liquid crystal conductive polymer strands such as thoseshown in FIG. 2 take the form of a “lawn” having first ends attached tomolecular recognition headgroups 18 and second ends attached toelectrode 14. As will be described below, headgroups 18 may take manydifferent forms. Generally, they should change physical or chemicalstate in response to the presence of a particular component in analyte20. In a preferred embodiment, molecular recognition headgroups 18 areenzymes which undergo a redox transformation in response to contact witha specified analyte component. For example, the analyte may include aligand or substrate component 25 which selectively binds with and ischemically modified by headgroups 18. Preferably, the chemicalmodification is accompanied by generation of electrons which candirectly transferred to strands 16 and from there to electrode 14.Depending upon the type of molecular recognition headgroup 18 employedin the sensor 12, the thickness of a headgroup layer on top of theconductive polymer lawn 16 may be between about 5 and 150 angstroms.

[0083] Importantly, no mediator is required in this sensor design, soelectron transfer is direct and fast from headgroup 18 to electrode 14.Further, because the polymer strands 16 are commonly oriented,headgroups 18 are optimally presented for sensing the desired analytecomponent. That is, headgroups 18 are not sterically hindered by polymerstrands 16 or other structures.

[0084] While the plurality of conductive polymer strands 16 may have arather uniform length as depicted in FIG. 2, this need not be the case.More frequently, the individual polymer strands will have a wide rangeof lengths. This will be due to inherent variations in polymerizationtechniques or the polymer shearing techniques. Of course, thedistribution of polymer strand lengths can be made more uniform bypassing a raw collection of polymer strands through a chromatographycolumn, electrophoretic gel, ultrafiltration membrane, or other sizingapparatus. In a preferred embodiment, the average strand length ofconductive polymer strand 16 is between about 2 and 1,000 Å. Morepreferably, the length is between about 10 and 100 Å, and mostpreferably between about 3 and 40 Å. When DNA is employed as theconductive strands, the width of the individual sensor strands is in theneighborhood of 20 Å.

[0085] In a preferred embodiment, the substrate 14 is a p-type electrodeof a silicon photodiode. It may include, though this is not alwaysnecessary, a metallic back plate 22 for providing an ohmic contactbetween polymer strands 16 and bulk silicon electrode 14. Such backmetal plates are conventionally used in semiconductor devices asterminals for connection to an external circuit. The back metal plate 22may be made from any suitable conductive metal or alloy, including butnot limited to aluminum, copper, silver, gold, and platinum. Region 24represents the close packed liquid crystal spacing between EMOLEdeposited molecular recognition headgroups. Molecular recognitionheadgroups whose dimensions are greater than the width of underlyingmolecular wires to which they are attached occupy region 24.

[0086] In a preferred embodiment, the semiconductor substrate forms partof a rectifying diode such as a photodiode. FIG. 3 provides a schematicillustration of a photodiode based biosensor in accordance with oneembodiment of the present invention. A sensor 50 includes a photodiode52 including an n-type region 53 and a p-type region 54. Generally, anyconventional photodiode may be employed with this invention, but itshould have a surface suitable for affixing conductive polymer strandsand molecular recognition headgroups as described above. To this end,p-type region 54 may be provided with or without a back metal ohmiccontact 56 as shown. A plurality of strands of conductive polymer 58 areaffixed at one end to back-metal plate 56. The other ends of polymerstrands 58 are attached to a collection of molecular headgroups 62. Theresulting structure, as illustrated, may be identical with the structureof elements 14, 22, 16 and 18 as shown in FIG. 2.

[0087] Photodiode 52 includes a depletion region 60 which automaticallyforms at the p-n semiconductor junction. As is known to those of skillin the art, depletion regions form at these interfaces because mobileholes diffuse from p-type regions into n-type regions just across theinterface where they are combined with electrons available in the n-typeregion. Similarly, mobile electrons in the n-type region diffuse acrossthe interface to the p-type region where they combine with holes. As aresult, within the reach of charge carrier diffusion, essentially allmobile charge carriers are depleted.

[0088] When light (or other radiant energy of appropriate wavelength) isshown on a photodiode such as photodiode 52, some holes and electronscross the semiconductor band gap and provide additional mobile chargecarriers which can be drawn out of photodiode 52 by an applied potentialor external short circuit connection. Applied potentials or externalshort circuit connections may be made through a digital multi-meter 64,a variable potential power supply, a battery, another photodiode, or apotentiostat, for example. Of course, many other potential sources orexternal short circuit connections may be employed. A multi-meter 64 hasthe advantage of being inexpensive yet able to detect the amount ofcurrent flowing as a result of the incident light. Additional electronsare attracted to p-type region 54 by the excess holes generated by thelight. Similarly, electrons flow out of n-type region 53 because thereare now excess electrons by virtue of the light excitation. This currentflows through a line 66, multi-meter 64, and a line 68. Note that line68 is electrically connected to back plate 56. Similarly, line 66 isconnected to a metal back plate 70.

[0089] When electrons are injected into the p-type region 54, they maycombine with and thereby annihilate holes. Thus, the photocurrentamplitude is reduced. Detection of this deviation from normalphotocurrent specifies that an analyte component has been detected. Ithas been found that the amplitude of this deviation is proportional tothe analyte component concentration. Further, it has been found that thedeviation is present in both the current and voltage associated with thephotodiode.

[0090] It should be understood that the sensors of this embodiment ofthe invention can be formed on any type of diode in which an externalstimulus generates a baseline current. Such stimulus may be heat(thermally generated charge carriers), electric field, radiation, etc.In each case the baseline current is at least partially “quenched” byelectrons or holes injected from the lawn of molecular devices when aspecified analyte component is present. Amplitude of the deviation frombaseline is often proportional to concentration of the analytecomponent. A simple calibration curve for each chip can be used todetermine concentration of the analyte component(s) in unknown samples.

[0091] In a particularly preferred embodiment, the sensor is dividedinto a plurality of regions, each capable of sensing the presence of adifferent analyte component. For example, a first region might include,as molecular recognition headgroup, glucose oxidase to sense thepresence of glucose, a second region might include cholesterol esteraseand cholesterol oxidase to sense the presence of cholesterol, a thirdregion might include alcohol dehydrogenase to sense the presence ofethanol, etc. Each of these regions will be separately addressable byelectronic circuitry to uniquely identify the presence a particularanalyte component. Each of the sensor regions could be made separatelyaddressable by specialized circuitry employed in conventional integratedcircuits. While the circuitry need not be particularly complex, suchdevices allow very sophisticated processing of the data provided by thesensor regions.

[0092] The molecular devices (headgroup and conductive strand affixed toan electrode surface) in each region may be formed by processes similarto those employed in integrated circuit fabrication. For example,certain regions could be exposed to light radiation shown through apatterned reticle. Those regions would be selectively activated orprotected depending upon the use of appropriate chemical protectinggroups. A liquid crystal conductive polymer region or headgroup regionwould then be formed on the reactive regions. Such processes aredescribed in U.S. Pat. No. 5,252,743 issued to Barrett et al. andPritchard et al., “Micron-Scale Patterning of Biological Molecules”Angew. Chem. Int. Ed. Engl., Vol. 34, No. 1, pages 91-93 (1995), forexample, which is incorporated herein by reference for all purposes.Alternatively, an electric potential could be selectively applied tocertain of the substrate regions to selectively electrodeposit thedistinct sensor regions.

[0093] II. Solid Substrate

[0094] Various solid substrates may be employed in the invention. Thesolid substrate should undergo a detectable change in response to anelectrical stimulus from the molecular wire. The substrate material maybe biological, nonbiological, organic, inorganic, or of a combination ofany of these, existing as particles, strands, precipitates, gels,sheets, tubing, spheres, containers, capillaries, pads, slices, films,plates, slides, etc. The substrate may have any convenient shape such asdisc, square, sphere, circle, etc. The substrate and its surfacepreferably, though not necessarily, form a rigid support on which tocarry out the reactions and fabrication processes described herein. Thesubstrate and its surface may also be chosen to provide appropriatecrystal or non-crystal lattice structure, wafer or thin filmorientation, n- and p-type doped materials, surface texture, back metalpattern, grid metal pattern, surface chemistry, etc. The raw macro-solidsubstrate may be composed of a semiconductor or standard electricalcomponent. Preparation of surfaces by lapping, polishing, chemicaltreatment, ion implantation, photolithography, etching, chemical vapordeposition (CVD), molecular beam epitaxy (MBE), etc. may provide apatterned macro-solid substrate suitable for further processing by meansof the present invention.

[0095] Various semiconductor substrates may be employed in theinvention. The semiconductor substrate may be biological (e.g., lipidbilayers, membrances, detergent solubilized membrane fragmentscontaining embedded protein electron transport pathways, blood brainbarrier (BBB), epithelial linings, intestinal linings, intracellularmembrane fragments, intracellular organelles, different tissue cellsurface types, membrance surfaces from different blood types of redblood cells, membrane surfaces from different types of lymphocytes,macrophages, and white blood cells, lyposomes, arterial and venous bloodvessel walls, neuronal conduction pathways, etc.), nonbiological,organic, inorganic, or of a combination of any of these. Usually, thesemiconductor substrate will be composed of silicon, doped diamond,indium tin oxide, tin oxide, gallium arsenide, cadmium sulfide, cadmiumselenide, cadmium telluride, germanium, copper indium diselenide, copperindium disulfide, copper indium ditelluride, zinc sulfide, zincselenide, mercury telluride, mercury selenide, graphite, etc. orcombinations thereof. Other substrate materials will be readily apparentto those of skill in the art upon review of this disclosure. In apreferred embodiment the semiconductor substrate is a p-n dopedpolycrystalline or monocrystalline silicon (e.g., having a surfacecrystallographic orientation in the <100> or <111> direction) or copperindium diselenide monocrystalline thin film deposited onto glass.

[0096] A semiconductor substrate may form part of a homojunction devicewhere the same semiconductor material is employed on either side of thep-n junction, differing only in dopant type; or heterojunction device,where the materials on either side of the p-n junction aresemiconductors but different semiconductors. Processes and chemistriesfor homo- and heterojunction device manufacture are known in the art andwill not be described in significant detail. A conventional photovoltaicsolar cell is an example of a semiconductor homojunction device. It is astandard n-p junction, rectifying diode with contact metallizationpartially covering its emitter to allow light entrance.

[0097] In a rectifying diode, for example, conducting back metal contactpatterns may be located on the p-type surface and conducting grid metalcontact patterns may be located on the n-type surface. Such back metalpatterns are generally used for the purpose of providing an ohmiccontact to the semiconductor diode. In the present invention, they maybe used for attaching highly conductive terminal contacts of theconducting polymer to the semiconductor substrate surface in specificregions as described in the next section. Back or grid metal contactsare typically made from a conductive metal layer such as aluminum,copper, gold, silver, etc. The back or grid metal may be textured andmay adopt lattice matching of underlying monocrystalline <100> or <111>silicon surfaces upon which it is deposited. Alternatively, theconducting polymers or thin films of this invention may be directlyconnected to p-type polycrystalline or monocrystalline surfaces, withoutthe need for back metal.

[0098] The raw macro-solid substrate may be connected to or comprisestandard electrical components (e.g., transistor, diode, electrode,semiconductor heterojunction, semiconductor homojunction, Schottkybarrier, capacitor, resistor, inductor, CMOS, TTL CMOS, FET, ISFET,MOSFET, ENFET, REFET) or combinations thereof (See e.g., U.S. Pat. No.5,126,921 to Fujishima et al.; U.S. Pat. No. 5,108,819 to Heller et al.;U.S. Pat. No. 5,403,700 to Heller et al.). Memory and logic circuitry onsuch chips can be employed to interpret sensor signals. In a preferredembodiment, the sensor wiring will be attached to transistor gates,sources, or drains (to control potential) or to other circuit or devicecomponents to control current. Preparation of active surfaces on thesemiconductor substrate may be accomplished by various fabricationtechniques including, for example, lapping, polishing, chemicaltreatment, ion implantation, photolithography, etching, chemical vapordeposition (CVD), molecular beam epitaxy (MBE), etc.

[0099] It may be possible to wire only few or even one conductivepolymer strand to a device element such as gate of a FET. Usingavailable technology reported by Yoo et al. in Science, entitled“Scanning Single-Electron Transistor Microscopy: Imaging IndividualCharges”, Vol. 276, pages 579-582 (1997) (which is incorporated hereinby reference for all purposes), source, drain, and gate elements of verysmall dimensions have been fabricated on a scanning tunnellingmicroscope (“STM”) tip. Such devices have been reported to detecttransfer of single charge carriers. By attaching one or a few conductivepolymers (and associated headgroups) to the gate of such device, forexample, a single binding event (at single headgroup) could be detected.If the individual devices are made separately addressable, each polymerstrand/headgroup combination could form a molecular transistor of verysmall dimensions. Separately addressable STM tips are discussed byService in Science, “Atomic Landscapes Beckon Chip Makers and Chemists”Vol. 274, pages 723-724 (1996).

[0100] III. Sequential Electrochemical and Chemical DepositionTechniques

[0101] Sequential electrochemical or chemical deposition techniques maybe used to attach molecular recognition surfaces to conductive polymersand to attach conductive polymers onto semiconductor wafer substratesprepared as described above. Specifically, the present process methodsof this invention may employ various processes related toelectrochemical atomic layer epitaxy (ECALE), sequential monolayerelectrodeposition (SMED), and thin film chemical deposition (CD) in aprocess referred to herein as electrochemical molecular layer epitaxy(EMOLE) to deposit, polymerize, and/or orient monomers, polymers,macromolecules, or thin films into liquid crystal conducting polymers or“molecular wires” with highly conductive terminal contacts. Preferably,one terminal contact of the formed one-dimensional molecular wire is“molecularly soldered” or electrically connected to the substratesurface (i.e., the back metal coated on a p-type surface of thesemiconductor homojunction substrate). The other terminal contact isdirected outward by virtue of extended liquid crystal conducting polymerorientation perpendicular to the substrate surface as illustrated abovein FIG. 2. Repeat of analogous deposition techniques are used to“molecularly solder” or electrically connect an active molecularrecognition headgroup to the free terminal contacts (also illustrated inFIG. 2) permitting rapid and direct charge conduction from the molecularrecognition sites to the semiconductor substrate.

[0102] In a preferred embodiment of the invention, sequential depositionoccurs only in specific regions of the semiconductor substrate (e.g., onspecific electrically or chemically activated surface regions of thesubstrate electrode). This provides a patterened surface of individuallywired molecular recognition sites.

[0103] Examples of three sequential deposition techniques(electrochemical and chemical) and their application to production ofatomic layers of compound semiconductors and conducting polymers aredescribed below in Section III, A-C. A modified form of these processescalled electrochemical molecular layer epitaxy (EMOLE) may be employedto fabricate a single sensor site or an array of sensor sites.

[0104] A. Electrochemical Atomic Layer Epitaxy (ECALE)

[0105] The epitaxial growth of semiconductors is an important and activearea of research. The development of new, low temperature techniques forthe preparation of high-quality semiconducting thin-film materials is offundamental importance to the semiconductor chip industry. Considerableeffort has been devoted to study the epitaxial growth of these materialsin vacuum (e.g., molecular beam epitaxy (MBE). Electrodepositionrepresents an alternative to the expense of vacuum techniques. Inaddition, electrochemistry is usually performed near room temperature,and therefore avoids the interdiffusion problems associated with thehigh temperatures used in vacuum deposition methods. Research has beendirected towards the epitaxial electrodeposition of II-VI compoundsemiconductors. A method for epitaxial electrodeposition and digitaletching, electrochemical atomic layer epitaxy (ECALE), is beingdeveloped. The method involves the alternated electrodeposition ofatomic layers of the constituent elements which make up a compound.Deposition is limited to an atomic layer by the use of underpotentialdeposition (UPD). UPD refers to a surface-limited process whereby adepositing element forms a compound with substrate surface atoms at apotential below that required for bulk deposition of the element.Deposition of the element proceeds until the surface is “covered”. Afterthe surface is covered, subsequent deposition requires a higherpotential to promote bulk deposition. Thus, UPD is usually limited tomonolayer coverage.

[0106] ALE (atomic layer epitaxy) refers to a series of vacuum basedmethods for semiconductor growth where a compound is formed a monolayerat a time by the alternated deposition of atomic layers of theconstituent elements. ALE is applicable to a variety of thin filmformation methods such as molecular-beam epitaxy (MBE), metalloorganicmolecular beam epitaxy (MOMBE), chemical vapor deposition (CVD),metalloorganic chemical vapor deposition (MOCVD), etc. These vacuummethods involve such problems as the need for careful control ofreactant fluxes in order to obtain epitaxial deposits. ALE is currentlyunder development which allows less stringent control of growthparameters. Unique to ALE is compound growth of one atomic layer at atime. This technique relies on surface-specific reactions which resultin only a monolayer of reactivity. If the reactant is an elementalvapor, the substance temperature is adjusted so that bulk depositssublime while the first monolayer remains due to an enhanced stabilityresulting from compound formation. After pumping (evacuation) of thefirst element, a similar procedure is performed with the second element.For a compound such as CdTe, a layer of Cd is formed followed by a layerof Te. Thin film growth is achieved by repeating the cycle.

[0107] In the formation of a compound such as GaAs by ALE in the MOCVDmode, a flux of H₃As, an arsenic precursor gas, is exposed to thesubstrate at a temperature which allows formation of a single As surfacelayer. All excess H₃As is subsequently pumped away under high vacuum.The As atomic layer is stabilized by compound formation with previouslydeposited Ga. A flux of tetramethyl gallium (TMG), a gallium precursorgas, is then exposed to the surface, and similarly an atomic layer of Gais formed. Excess gas is pumped away under high vacuum. Thin films areproduced by repeating this cycle.

[0108] ECALE is the electrochemical analog of atomic layer epitaxy (ALE)employing UDP in place of temperature control to deposit monolayers. Useof UPD in order to electrodeposit atomic layers of both elements, atpresent, requires that one element be deposited by reductive UPD whilethe other is deposited by oxidative UPD. In this way, oneunderpotentially deposited element can be held on the surface at thepotential used subsequently to deposit the other element. In theformation of a compound such as CdTe, Te can be oxidativelyunderpotentially deposited from Te²⁻ at a fairly negative potential.Cadmium can next be reductively underpotentially deposited from a Cd²⁺solution at a more positive potential, where previously deposited Teremains stable. Electrodeposited semiconductors do not have to beannealed as in ALE which is typically done for 15 minutes at 300° C.

[0109] Digital etching, the reverse process of deposition, is a naturalextension of the ECALE method. Increasing the negative voltage potentialto strip or etch monolayers is possible. A method for the digitalelectrochemical etching of compound semiconductors in an electrochemicalflow cell system in which alternating electrochemical potentials areapplied between a reference electrode and the compound semiconductorsufficient to strip portions, preferably atomic layers, of the elementsof compound semiconductors from the compound semiconductors is describedin Stickney et al.: U.S. Pat. No. 5,385,651 and Stickney et al.: WO94/28203.

[0110] B. Sequential Monolayer Electrodeposition (SMED)

[0111] Sequential Monolayer Electrodeposition (SMED) provides monolayersof II-VI compound semiconductors and is related to the ECALE methoddescribed above. However, unlike the ECALE method, all depositedelements are provided in the same electroplating solution. They arecodeposited and then one which deposited in excess is electrochemicallystripped away. For example, Cd²⁺ and Se²⁻ may be deposited from the sameelectroplating solution by cyclic voltammetric deposition at fast scanrates with a nickel rotating disk electrode. The procedure was designedto eliminate the problem of bulk Se formation, using a cyclic depositionscheme that cathodically deposits submonolayer amounts of CdSe and alarge stoichiometric excess of Cd. The excess Cd is then stripped off bysweeping the electrode to a positive potential as part of thevoltammetry cycle (Cd is readily stripped close to its thermodynamicreduction potential). Since the CdSe phase has a large negative freeenergy of formation (ΔG^(o) _(f,298K)=−141.5 kJ mol⁻¹), it was thoughtthat any free Se that is deposited in this process will react with theexcess Cd to form CdSe and not lead to large amounts of excess Se in thefilm. The net result is thus the sequential deposition of stoichiometricCdSe a monolayer (or less) at a time. It has been reported that such aprocedure leads to compositionally homogeneous, stoichiometric films andmay be a general method to electrodeposit binary materials with largethermodynamic or kinetic stabilities. (Kressin, A M; Doan, V V; Klein, JD; Sailor, M J: “Synthesis of Stoichiometric Cadmium Selenide Films ViaSequential Monolayer Electrodeposition” Chem. Mater. 3(6): 1015-1020,1991).

[0112] C. Thin Film Chemical Deposition (CD)

[0113] Conducting polymers continue to look promising as the activeelements of electronic and chemical devices such as flexiblelight-emitting diodes, chemical sensors and photovoltaic devices. As aresult, the thin film processing techniques for these materials havebecome increasingly important to the successful fabrication andoptimization of useful all-organic thin film devices. Techniques such asspin coating, electrochemical deposition, and Langmuir-Blodgett thinfilm transfer have all been utilized with varying degrees of success tomanipulate conjugated polymers into thin films. Fou et al. (Fou, A C;Ellis, D L; Rubner, M F: Molecular-Level Control in the Deposition ofUltrathin Films of Highly Conductive, In-Situ Polymerized P-DopedConjugated Polymers. Mater. Res. Soc. Symp. Proc. 328:113-118, 1994.)has described a thin film processing technique that has been developedfor the fabrication of ultrathin films of conducting polymers withangstrom-level control over thickness and multilayer architecture.Molecular self-assembly of in-situ polymerized conjugated polymersconsists of a layer-by-layer process in which a substrate is alternatelydipped into a solution of a p-doped conducting polymer (e.g.,polypyrrole, polyaniline) and a solution of a polyanion. In-situoxidative polymerization produces the more highly conductive,underivatized form of the conjugated polymer, which is deposited in asingle layer of precisely controlled thickness (30 to 60 Å). Thethickness of each layer can be fine-tuned by adjusting the dipping timeand the solution chemistry. The surface chemistry of the substrate(i.e., hydrophobic, charged, etc.) also strongly influences thedeposition, thereby making it possible to selectively deposit conductingpolypyrrole onto well defined regions of the substrates.

[0114] D. Electrochemical Molecular Layer Epitaxy (EMOLE)

[0115] Electrochemical molecular layer epitaxy (EMOLE) is a processingtechnology used to engineer the structure and properties ofmacromolecules deposited on a substrate surface in order to producehighly organized molecular materials. Preferably, this processing yieldsliquid crystal structures of the type described above. Typically,crystallization is viewed as producing homogenous and well orderedmaterials made of one or a few kinds of atoms or small molecules. It isalso possible though to crystallize larger and more complex moleculessuch as proteins, DNA, supramolecular assemblies such as ribosomes, andeven virus particles with atomic masses in excess of 100 milliondaltons. In fact, this is a necessary step in elucidating the structureof many macromolecules. Co-crystallization of two or more differentcomponents is also possible. The present invention provides EMOLEtechniques to produce layers of two-dimensional crystals or generallywell ordered arrangements of interconnected macromolecules for theproduction of a biosensor. EMOLE as described herein generally employslow current density and potential (which maintains the Helmholtzdouble-layer) to deposit uniaxially oriented liquid crystal conductingbiopolymers (proteins and DNA) at substrate surfaces.

[0116] Preferably, methods of this invention employ EMOLE to deposit,attach, polymerize, and/or orient monomers, polymers, macromolecules, orthin films into liquid crystal conducting polymers or “molecular wires”with conductive terminal contacts. “Thin film” is a term used herein tomean a well defined atomic or molecular deposition layer on a flattwo-dimensional substrate. Thin films can be made by many techniques(i.e., ALE, CVD, Langmuir-Blodgett, dip coating, spin coating, EMOLE,etc.) and be composed of many materials. Thin films can sometimes becharacterized as a “lawn” or “liquid crystal.”

[0117] Conditions which promote oriented liquid crystal polymers will bepresented below. EMOLE may be employed to form conductive electronicconnections at each end of the oriented liquid crystal conductingpolymers (i.e., the headgroup end and the substrate end). By connectingthem at a first end of conductive polymer strands in a liquid crystalorientation, the molecular recognition headgroups are stericallyunhindered in their chemical or biochemical binding/recognition ofanalyte species. As a consequence of an analyte binding event to amolecular recognition site, rapid electron or hole transfer from theoriented liquid crystal molecular recognition site through the attachedoriented liquid crystal conducting polymer or thin film, to thesemiconductor substrate will produce a signal. The amplitude of thesignal or number of electrons or holes tunneling to the semiconductorsurface taken in aggregate will reflect the amount of specified analytespecies present.

[0118] In a preferred embodiment of this invention, a firstelectrodeposition cycle affixes strands of a conducting polymer on asubstrate (e.g., a p-type surface of a semiconductor such as a p-njunction solar cell described above). This is depicted in FIG. 4A. Inthis cycle, a first medium 402 containing a polymer 404 to be deposited(or a precursor of that polymer such as monomers) is contacted with asubstrate 406. Preferably, though not necessarily, medium 402 is aliquid solubilizing the polymer strands. Medium 402 may be held within acontainer 407 as shown, or may passed over substrate 406 in a continuousflow reactor. A potential is then applied to substrate 406 via a circuit408 to drive the first cycle and deposit a lawn of immobilized polymerstrands 410. Note that circuit 408 includes substrate 406, medium 402, acounter electrode 412, and a power supply 414. If polymer strands 404have a positive charge, then a negative potential is applied to thesubstrate; but if they have a negative charge, a positive charge isapplied to the substrate. In either event, the potential and/or currentdensity should be controlled to ensure that (1) the polymer is affixedto the substrate with strength to allow electron transport, and (2) thedeposited polymer strands have a substantially common orientation. Itmay be desirable to include a charge group on only one end of polymers404 so that that end is selectively coupled to the surface of substrate406. If the polymer strand is a nucleic acid, the charge group could beattached by including it at one end of a nucleic acid strand (designedmuch like a conventional nucleic acid probe) which strand iscomplementary to an end of the nucleic acid to be affixed. Of course,other techniques for attaching charge groups (or other functionalgroups) to one end of a polymer strand are known in the art and may beprofitably employed in the context of the present invention.

[0119] In a specific embodiment, electrodeposition current densityranging from about 10 to 300 μA cm⁻² and voltage potential ranging fromabout 10 to 300 mV can be generated by light induced photoconduction atthe n-type and p-type surfaces of the submerged solar cell. Depositioncycle variables include i) applied potentials (i.e., magnetic/voltage);ii) solution condition (i.e., concentration of deposited material, pH,electrolyte, solvent, temperature, etc.); and iii) semiconductorsubstrate (i.e., polycrystalline, monocrystalline, single-crystal faceorientation, smooth or textured surface, metal contact coating, latticematching of coating, etc.). As will be understood to those of skill inthe art, these variables may be adjusted to produce an optimalmolecular-scale structure.

[0120] For example, the following guidelines may be employed to depositsuitable molecular wires. First, applied potentials must be low enough(e.g., 0.001 to 1500 mV) to maintain a Helmholtz double-layer duringelectrodeposition of conducting polymers and molecular recognitionheadgroups onto semiconductor substrate. Applied potential ranges willvary depending on the size, charge density, counter ion, and viscosityof the to be deposited material. Second, current densities must be lowenough (e.g., 0.001 to 1500 μA cm⁻²) to maintain a Helmholtzdouble-layer during electrodeposition of conducting polymers andmolecular recognition headgroups onto semiconductor substrate. Currentdensity ranges will vary depending on the size, charge density, counterion, and viscosity of the to be deposited material. Third, thesemiconductor substrate should be chosen to maintain a Helmholtzdouble-layer during electrodeposition of a uniaxially oriented liquidcrystal structure on the surface of the semiconductor substrate. Asnoted, it may be polycrystalline or monocrystalline, having smooth ortextured surface. It may also have a metal contact coating.

[0121] Further, the solution conditions should meet certain specificcriteria. For example, the concentration of deposited material should below enough (e.g., 0.001 to 10 mg/mL) to maintain a Helmholtzdouble-layer during electrodeposition of conducting polymers andmolecular recognition headgroups onto semiconductor substrate. Further,the pH should be adjusted to approximately two (2) pH units above orbelow the pK_(a) or pI of the conducting polymer or molecularrecognition headgroup to produce a polymer of opposite charge from thesurface of the semiconductor substrate. Still further, the electrolyteshould be chosen to have a counter ion type and electrolyteconcentration (e.g., 0 to 150 mM salt) that maintains a Helmholtzdouble-layer during electrodeposition of a uniaxially oriented liquidcrystal structure on the surface of the semiconductor substrate. Highelectrolyte concentration will produce too much current and destroy theHelmholtz double-layer during electrochemical deposition processing. Inaddition, the solvent should be chosen from a range of organic andaqueous solvents and co-solvents to maintain a Helmholtz double-layerduring electrodeposition of a uniaxially oriented liquid crystalstructure on the surface of the semiconductor substrate. Conductingpolymers and molecular recognition headgroups should be soluble in thesolvent or co-solvent used. Finally, the temperature should be greaterthan the freezing point (fp) and less than the boiling point (bp) of thesolvent or co-solvent to maintain a Helmholtz double-layer duringelectrodeposition of a uniaxially oriented liquid crystal structure onthe surface of the semiconductor substrate.

[0122] During the sensor formation process, a second electrodepositioncycle is performed to attach molecular recognition sites on top of theunderlying uniaxially oriented liquid crystal conducting polymer layer.The second cycle is depicted in FIG. 4B. As with the first depositioncycle, a desired material is deposited from a medium; preferably aliquid medium 422. In this case, second medium 422 contains headgroups420, or an appropriate precursor, to be deposited. After medium 422 isbrought into contact with substrate 406 (to which polymer strands 410were affixed in the first cycle), a potential is applied to thesubstrate through circuit 408 to drive the second cycle. The potentialwill be positive or negative depending upon the charge on theheadgroups. This results in deposition of a lawn of immobilizedheadgroups 424 attached to an unfixed end of polymer strands 410. Thepotential and/or current density should be controlled to ensure that (1)the headgroup is affixed to the polymer strands with strength to allowelectron transport, and (2) the deposited headgroups have asubstantially common orientation. Deposition cycle variables areadjusted to ensure production of a single molecular layer of uniaxiallyoriented liquid crystal chemically or biologically active molecularrecognition sites 424 individually “wired” to underlying uniaxiallyoriented liquid crystal electrically conducting polymer layer 410. Theheadgroups to be deposited may be provided with one or more functionalgroups which direct the headgroups onto strands 410 in a desiredorientation. As with the polymer strands, the headgroups may befunctionalized with a charge group. In many cases, it may be desirableto locate the charge group away from the active site of the headgroup,so that the headgroup will attach with the active site facing themedium.

[0123] Deposition conditions must be tailored to the material to bedeposited. In one embodiment of this invention, DNA deposition and GODenzyme deposition conditions happened to use similar current density andapplied potential (e.g., 10 to 300 μA cm⁻² and 10 to 300 mV). However,solution conditions in the two deposition cycles (i.e., concentration ofdeposited material, pH, electrolyte, solvent) will not be the same.

[0124] As should be apparent, the deposition reactions require that thepolymer strands and recognition headgroups be electrically charged andmobile in an electric field. Thus, the compositions of the first andsecond media may have to be carefully chosen. Typically, though notnecessarily, the first medium is removed and the substrate is allowed todry before being contacted with the second medium.

[0125] 1. Deposition of Uniaxially Oriented Liquid Crystal ConductingBiopolymers (Proteins and DNA)

[0126] In a preferred embodiment of the present invention, EMOLE methodsare used to sequentially deposit, attach, and orient liquid crystalconducting polymers (e.g., DNA and proteins) onto the surface of asubstrate (e.g., a p-type silicon of a polycrystalline p-n junctionsolar cell). For example, the pH of a DNA-electrolyte depositionsolution is adjusted to ˜6.0 (more than two pH units above the pK_(a) orpI of DNA) producing a negatively charged DNA biopolymer. Light inducedphotoconduction by a submerged solar cell generates an electric field inthe DNA-electrolyte solution which uniaxially orients negatively chargedDNA strands onto the positive p-type silicon surface. Solar cell appliedcurrent density and voltage potential are low enough to establish andmaintain a Helmholtz double-layer (as described in U.S. Pat. No.5,215,631 to Westfall) between the p-type silicon surface and the DNAand counter ions in solution. The very gentle EMOLE conditionsfacilitate electrochemical deposition of uniaxially oriented liquidcrystalline extended DNA structures orthogonal to the semiconductorsubstrate surface. By “gentle,” it is meant that the conditions preservethe Helmholtz double-layer as described in the Westfall referencediscussed above.

[0127] EMOLE methods may be used to sequentially deposit, attach, andorient liquid crystal conducting protein (i.e., molecular recognitionsites) on top of the underlying uniaxially oriented liquid crystal DNAlayer affixed to the surface of the silicon substrate chip. For example,the pH of a protein-electrolyte deposition solution is ˜7.0 (more thantwo pH units above the pK_(a) or pI of the protein) producing anegatively charged protein biopolymer. Light induced photoconduction bya submerged solar cell generates an electric field in theprotein-electrolyte solution which uniaxially orients negatively chargedproteins onto the “lawn” of liquid crystal DNA molecular wires. Solarcell applied current density and voltage potential are low enough toestablish and maintain a Helmholtz double-layer between the DNA-modifiedp-type silicon surface and the protein and counter ions in solution. Thevery gentle EMOLE conditions facilitate sequential electrochemicaldepositions that maintain the first monolayer of uniaxially orientedliquid crystalline extended DNA structures orthogonal to thesemiconductor substrate surface while depositing a second monolayer ofuniaxially oriented liquid crystalline protein “headgroups” on top ofthe underlying “lawn” of liquid crystal DNA wires as characterized bythe following references: Collings, P J: Chap. 3. Electric and MagneticField Effects. In: Liquid Crystals: Nature's Delicate Phase of Matter.Princeton University Press; Princeton, N.J.; 1990; pp. 35-55. Collings,P J: Chap. 9. Polymer Liquid Crystals. In: Liquid Crystals: Nature'sDelicate Phase of Matter. Princeton University Press; Princeton, N.J.;1990; pp. 162-180. Pelzl, G: Chap. 2. Thermodynamic Behavior andPhysical Properties of Thermotropic Liquid Crystals. In: LiquidCrystals. Stegemeyer, H; guest ed. Steinkopff, Darmstadt and Springer,New York; 1994; pp. 51-102. Zentel, R: Chap. 3. Liquid CrystallinePolymers. In: Liquid Crystals. Stegemeyer, H; guest ed. Steinkopff,Darmstadt and Springer, New York; 1994; pp. 103-141).

[0128] Upon electrochemical deposition of a monolayer of uniaxiallyoriented liquid crystal protein, the DNA-silicon substrate is removedfrom the deposition bath and allowed to slowly dry and cool in thepresence of an applied electric field. This allows the oriented liquidcrystal protein structure to be “locked-in” on top of the orientedliquid crystal DNA molecular wire terminal surface of the dry siliconsubstrate chip as described in the following references: Collings, P J:Chap. 6. Liquid Crystal Displays. In: Liquid Crystals: Nature's DelicatePhase of Matter. Princeton University Press; Princeton, N.J.; 1990; pp.96-120. Albrecht, C; Enkelmann, V; Lieser, G; Schwiegk, S; Wang, W;Wegner, G; Zierer, D: The Crystallization Behavior of Rod-LikeMacromolecules. In: Crystallization of Polymers. Dosiere, M; ed. KluwerAcademic Publishers; Dordrecht, Boston, London; 1993; pp. 323-330.Brandes, R: Part I. Generation of Tailored Radio Frequency Pulses ForNMR. Part II. Deuterium NMR Studies of Oriented DNA, and Its InteractionWith Water. Dissertation, Ph.D. in Chemistry; University of California,San Diego; 1988.

[0129] Because EMOLE employs an electrodeposition mechanism, the speciesto be deposited must be charged. Such charge exists naturally on manymaterials of interest when in the solution phase. However, manymaterials must be charged to facilitate EMOLE deposition. Manybiopolymers, for example, can be positively charged by adjusting the pHof the biopolymer-electrolyte deposition solution to more than two pHunits below the pK_(a) or pI of the biopolymer. The resulting positivelycharged species is suitable for electrochemical deposition onto negativen-type semiconductor surfaces, for example.

[0130] Like all liquid crystals, the oriented polymers of this inventionmay have their properties tailored by adding suitably functionalizedgroups of atoms to the polymer backbone. Such properties includemechanical strength as well as ferroelectricity, non-linear opticalactivity, and electronic charge transfer. The physical principlesinvolved are summarized in a number of books (Collings, P J: LiquidCrystals. Nature's Delicate Phase Of Matter. Princeton University Press;Princeton, N.J.; 1990. Stegemeyer, H (guest ed.): Liquid Crystals.Steinkopff, Darmstadt and Springer, New York; 1994. Plate, N A (ed.):Liquid-Crystal Polymers. Plenum Press; New York, London; 1993. Dosiere,M (ed.): Crystallization of Polymers. Kluwer Academic Publishers;Dordrecht, Boston, London; 1993). Anisotropic chemical and physicalproperties of liquid crystals and liquid crystal polymers are a resultof the molecular-scale structure formed. It was recently realized thatmanipulation of molecular-scale structure, and therefore function ofliquid crystals and liquid crystal polymers, not only depended on theuse of different functionalized organic molecules, but was heavilydependent on variables such as solvent, electrolytes, impurities,dopants; liquid crystal field effects (i.e., applied electric, magnetic,temperature, mechanical, electromagnetic radiation, or chemical fields);and processing techniques used (Collings, P J: Liquid Crystals. Nature'sDelicate Phase Of Matter. Princeton University Press; Princeton, N.J.;1990. Stegemeyer, H (guest ed.): Liquid Crystals. Steinkopff, Darmstadtand Springer, New York; 1994. Plate, N A (ed.): Liquid-Crystal Polymers.Plenum Press; New York, London; 1993. Dosiere, M (ed.): Crystallizationof Polymers. Kluwer Academic Publishers; Dordrecht, Boston, London;1993. Collyer, A A (ed.): Liquid Crystal Polymers: From Structures ToApplications. Elsevier Applied Science; London, New York; 1992. Lam, L;Prost, J (eds.): Solitons In Liquid Crystals. Springer-Verlag; New York,Berlin, Heidelberg, London; 1992). For example, coupling of molecularrecognition surfaces to electronically conducting polymers may resultfrom chiral smectic (layered cholesteric) liquid crystal structuresformed by sequential deposition of DNA and protein using EMOLEfabrication techniques provided by this invention. In a preferredembodiment, biopolymers (DNA and protein) and EMOLE techniques are usedto fabricate a molecular recognition (MR) device.

[0131] IV. Conducting Polymers and Thin Films

[0132] Many different conducting polymers and thin films can be employedfor “wiring” molecular recognition sites to a semiconductor or standardelectrical component substrate. Generally such polymers may bebiological, organic, inorganic, water soluble, lipid soluble orcombinations thereof. Many examples of conducting polymers suitable forthis invention are discussed by Skotheim, T A: Handbook Of ConductingPolymers. Vol. 1-2. Skotheim, T A, ed. Marcel Dekker, Inc.; New York,Basel; 1986. Types of conducting polymers and thin films suitable foruse in this invention include, but are in no way limited to thefollowing general classes: aromatic metal-doped polymers (e.g.,polyaniline doped by metal salts), π-stacked (aromatic) polymers (e.g.,polyphenanthroline; pyrazine-bridged polymers of π-stackedmetalloporphyrins; 2, 3, 6, 7, 10, 11-hexahexylthiotriphenylene (HHTT)),π-stacked (aromatic) helical polymers (e.g., DNA), organic π-conjugatedlinear polymers (e.g., polyacetylene), heterocyclic polymers (e.g. DNA,polyporphyrins), macrocyclic polymers (e.g., polyporphyrins with a redoxmetal; polytetrazacyclododecane with a redox metal), porphyrin polymers,polymer composites (e.g., layered polymer mixtures), polyelectrolytepolymers, (e.g., proteins, DNA), liquid-crystal polymers (e.g., certainproteins; DNA; polyporphyrins; and 2, 3, 6, 7, 10,11-hexahexylthiotriphenylene (HHTT)), self-organizing polymers (e.g.,polysurfactants with redox metal; HHTT), branched polymers, dendriticpolymers (e.g., starburst dendrimers with redox metal), chaotic polymers(e.g., poly (SiO₂)_(n) in glass with redox polymer), biopolymers (e.g.,protein, DNA, polyporphyrins), inorganic polymers (e.g. iron (hydrous)oxides), organometallic polymers (e.g., ferrocene polymers),inorganic/organic hybrid polymers (e.g. iron (hydrous)oxide/polybipyridine complex), metallocene polymers (e.g.,polyferrocene), inclusion compound polymers (e.g., polyzeolite withredox metal), mixed doped polymers, colloidal/sol-gel doped polymers(e.g., poly (SiO₂)_(n) with redox metal), ionomers (e.g., DNA, certainproteins, and certain polysurfactants), metal cluster doped polymers(e.g., iron (hydrous) oxide/polybipyridine complex), redox polymers(e.g., Heller (Osmium-PVP) and Skothiem (ferrocene-polysiloxane)), blockpolymers, graft polymers, transition metal films (e.g., deposited byatomic layer epitaxy (ALE)), high temperature superconductor films(e.g., atomic layer epitaxy (ALE) of appropriate redox metals),Langmuir-Blodgett films (e.g., detergents, amphiphiles, surfactants),sol-gel glass films (e.g., spin glass films), etc., or any combinationsof the above. Conducting polymers of appropriate strand lengths for eachof these may be employed herein.

[0133] In some cases, the native form of the polymer will be aninsulator, but upon appropriate doping, addition of impurities,hydration, conformational change, ionization, oxidation, reduction, etc.they become conductive. Further, some conducting polymers may bereversibly switched between conductive and insulative states.Polyaniline, for example, will become conductive in the protonated oroxidized form. Other “switchable” conductive polymers include, forexample, polymers polymerized from the following monomers:N-methylpyrrole, thiophene, 3-methylthiophene, 3,4-dimethylthiophene,vinylferrocene, styrene, nitrostyrene, viologens, vinyl-pyridine,vinyl-2,2′-bipyridine, vinylrubrene, quinone-based compounds, andderivatives thereof. This invention may also take advantage of suchconductivity transformation as a primary or auxiliary sensing mechanism.For example, a sensor signal may only be triggered by a combination oftwo events: a ligand binding with a molecular recognition headgroup anda pH change which causes the polymer wiring to become conductive.

[0134] Enzymes used in organic synthesis (i.e., to produce drugs andpharmaceuticals), may be used as molecular recognition headgroups ofthis invention. These include, but are not limited to, combinatorial andcommercial libraries of esterases, lipases, amidases, acylases, andother thermophillic and mesophilic enzymes with broad substratespecificities that can catalyze reactions in organic solvents and athigh temperatures. Upon ligand binding to an esterase or lipase, areaction will take place producing an alcohol and carboxylic acid fromthe cleaved ester bond. This will make the pH of theheadgroup/switchable polymer molecular environment more acidic; thus,protonating a reversibly switchable polymer to the protonated orconducting form. Amidase or acylase cleavage of an amide bond willproduce a free amine and a carboxylic acid. Chelation of the acid byanion exchange support would leave an increasing concentration of freeamine which would make the pH of the headgroup/switchable polymermolecular environment more basic; thus, deprotonating the reversiblyswitchable polymer to the neutral or insulating form. Examples ofenzymes used in organic syntheses may be used as molecular recognitionheadgroups to monitor levels of drugs and pharmaceuticals in the humanblood.

[0135] Esterase, lipases, acylases, or amidases may also be used todeprotect ligands to alcohols, carboxylic acids, or free amines whichthen become substrates suitable for a second molecular recognitionheadgroup, used to produce a signal by methods described in the presentinvention. For example, cholesterol esterase cleaves cholesterol esterfound in blood to cholesterol, which is then a substrate for cholesteroloxidase. Cholesterol oxidase would produce a signal much like glucoseoxidase described as an example of this invention.

[0136] Other approaches include, for example, Swager, et al. (Swager, TM; Marsella, M J; Conducting Polymers With Chemical Sensitive Traps andBarriers: New Molecule-Based Sensors. Mat. Res. Soc. Symp. Proc.328:263-266, 1994) which describes reversibly switchable polythiophenederivatives which exhibit large changes in bandgap in the presence ofspecific ions. These materials are based upon novel crown etherscontaining bithiophene monomers. Sensory polymers which are selectivefor K⁺ and Na⁺ are described. In such materials, specific ions induce atwisting of the polymers backbone, resulting in a decrease of π-orbitaloverlap between thiophene rings; reducing the extent of conjugationgiving rise to an insulating (higher bandgap) form.

[0137] Another example is of a sequence-specific DNA sensor. A specificsequence of single-strand DNA (nonconducting or insulating form) with 5′or 3′ terminus thiol could be adsorbed to a gold electrode substrate. Ananalyte sample containing the complementary DNA sequence would produce aDNA double-strand polymer which is a conducting form of DNA. This resultis a DNA sequence detector. DNA of the wrong sequence would not produceDNA double-strand polymer (conducting form). Appropriate end groupfunctionalities on single-strand DNA or no end group modifications ofsingle-stranded DNA (i.e., native DNA) using EMOLE methods could be usedto put sequence-specific single-strand (insulating form) DNA onsemiconductor substrates for use as a DNA sequence detector. DNA atcrime scenes could be identified on the spot, doing away with PCRtechniques and laborious and very costly DNA sequencing laboratoryprocedures.

[0138] Chemical-, photo-, or electro-polymerization of monomers may takeplace directly on the semiconductor or standard electrical componentsubstrate surface or pre-polymerized polymers may be deposited.Furthermore, once attached and polymerized, the polymer or thin film maybe oriented into a highly conductive liquid crystal polymer or thin filmform. This may be accomplished by depositing polymers in the presence ofappropriate electrical, magnetic, or chemical (solvent) fields.Preprocessing or conditioning of polymers is described in the Handbookof Polymer Synthesis (Plastics Engineering Series, Volume 24)Kricheldorf, H. F., 1991. Chemical polymerization may employ, forexample, H₂O₂, organoperoxides, or 2, 2′-azobisisobutyronitrile (AIBN).Photopolymerization may employ photons which generate photochemicalradicals which can initiate and propagate polymerization.Electropolymerization is currently employed to synthesize conductingpolymers.

[0139] A. Electron Transport Proteins

[0140] An example of a conducting biopolymer that may be useful for thisinvention is the electron transport protein. Electron transport proteinsare a product of millions of years of biological evolution, fine tuningthe function of electronic conduction. In nature, electron transportproteins often reside in, and are oriented by, a liquid crystallinelipid bilayer membrane. In this invention, the electron transportprotein may be deposited into a close-packed oriented two-dimensionalcrystalline structure by EMOLE crystallization processing techniques.This produces a surface structure suitably oriented as a plurality ofmolecular wire interconnects.

[0141] Proper deposition and orientation of proteins can be accomplishedby manipulation of the physical and chemical conditions duringcrystallization. The EMOLE technique allows a systematic approachunderstanding and optimizing the relevant parameters for depositingprotein or peptide polymers as wires for sensors. More generally, thenewly developed techniques of EMOLE provide for experimental control ofprotein crystal structure and function.

[0142] Electron transport proteins are in some embodiments suitable foruse with this invention because they perform some of the functiondesired for molecular electronic device (MED) fabrication—i.e., electronstorage and transfer at the molecular-scale. These properties arise fromthe alpha-helical and beta-pleated sheet structures of these biologicalmacromolecules and from their non-protein prosthetic groups. Theseprosthetic groups are inorganic-, organometallic-, or metal atomcofactors which are integral to the structure of protein. A particularlyinteresting protein is cytochrome b₅₆₂ of E. coli. This protein is small(12,000 daltons), has a single polypeptide chain folded into a simple4-alpha-helical motif, the x-ray structure is known to 2.5 Å, and mostimportantly, the single heme group is non-covalently bound. This lastproperty allows for the substitution of other porphyrin analogs with avariety of coordinated metal atoms, greatly increasing the experimentalflexibility of the system (Ulmer, K M: Chap. 29. Self-Organizing ProteinMonolayers As Substrates For Molecular Device Fabrication. In: MolecularElectronic Devices II. Carter, F L; ed. Marcel Dekker, Inc.; New York,Basel; 1987; pp. 573-590).

[0143] Photosynthetic electron transport proteins electronicallyconnecting photosystem II and photosystem I in plants, and mitochondrialrespiratory electron transport proteins are examples of conductingbiopolymer proteins oriented by a liquid crystalline lipid bilayermembrane—the chloroplast membrane (Clayton, R K: Light and LivingMatter, Volume 2: The Biological Part. McGraw-Hill Book Company, NewYork, 1971) and mitochondrial membrane; facilitating an extremelyefficient electron transfer chain via electron tunneling mechanism(Pethig, R: Chap. 9. Electronic Properties of Biomacromolecules. In:Dielectric and Electronic Properties of Biological Materials. John Wiley& Sons; Chichester, New York; 1979; pp.290-356).

[0144] Electron transport proteins that may be found among the proteinsparticipating in the respiratory chain of mitochondria are for example:flavoproteins, nonheme iron proteins, and cytochromes b, c₁, c, a, anda₃. With the exception of the electron donor, NADH, all of these areelectron transport proteins, shuttling two electrons from each moleculeof NADH to reduce ½ O₂ to H₂O. This downstream free energy electrontransport to O₂ is coupled to phosphorylative production of ATP, abiochemical energy currency.

[0145] Electron transport from photosystem II to photosystem I in thechloroplast membrane of green plants involves the electron transportproteins cytochrome b₅₅₉ or b₃ and cytochrome f. Electron transport fromphotosystem I involves the electron transport proteins ferredoxin andcytochrome b₆.

[0146] All of these electron transport proteins are juxtaposed to eachother in membranes with increasing standard oxidation-reductionpotentials facilitating a downward free energy transfer of two electronsfrom one electron transporting protein to the next in a highly orderedchain.

[0147] B. DNA Quantum Wires

[0148] A second example of a conducting biopolymer not normally thoughtof as electrically conductive until recently is DNA (Meade, T J andKayyem, J F: Electron Transfer Through DNA: Site-Specific Modificationof Duplex DNA with Ruthenium Donors and Acceptors. Angew. Chem. Int. Ed.Engl. 34(3):352-354, 1995. Murphy, C J; Arkin, M R; Jenkins, Y; Ghatlia,N D; Bossmann, S H; Turro, N J; Barton, J K: Long-Range PhotoinducedElectron Transfer Through a DNA Helix. Science 262:1025-1029, 1993.Meade, T J: Chap. 13. Electron Transfer Reactions Through the DNA DoubleHelix. In: Metal Ions In Biological Systems. Vol. 32. Interactions ofMetal Ions With Nucleotides, Nucleic Acids, and Their Constituents.Sigel, A; Sigel, H; eds. Marcel Dekker, Inc.; New York, Basel, HongKong; 1996; pp. 453-478. Stemp, E D A; Barton, J K: Chap. 11. ElectronTransfer Between Metal Complexes Bound To DNA: Is DNA A Wire? In: MetalIons In Biological Systems. Vol. 33. Probing of Nucleic Acids by MetalIon Complexes of Small Molecules. Sigel, A; Sigel, H; eds. MarcelDekker, Inc.; New York, Basel, Hong Kong; 1996; pp. 325-365. Arkin, M R;Stemp, E D A; Holmlin, R E; Barton, J K; Hormann, A; Olson, E J C;Barbara, P F: Rates of DNA-Mediated Electron Transfer BetweenMetallointercalators. Science 273:475-480, 1996). DNA is a biopolymerwith known solution and solid-crystal structures. In this invention,deposition of an oriented extended liquid crystalline DNA structureorthogonal to a solid-substrate surface may be achieved by EMOLEcrystallization processing techniques. This produces a surface structuresuitably oriented as a plurality of molecular wire interconnects.

[0149] While not wishing to be bound by theory, the following discussionis presented to illustrate the state of the art as to DNA as aconducting medium. There is still no consensus in the art as to whetherDNA can actually act as a wire. The debate is set forth generally byWilson (Wilson, D N A: Insulator or Wire, Chem. & Eng. News, 1997:33,Feb. 24, 1997). While such debate rages, the following discussionassumes that DNA is in fact a very good conducting polymer and is apreferred wire for use with the sensors and EMOLE methods of thisinvention.

[0150] Long distance electron movement through DNA (i.e., ˜40 Å or ˜12base pairs) has been confirmed only in experiments in a water solution.DNA has to be fixed to a terminal base, substrate, etc., coupled withthe controlling of the thickness and orientation of molecules in orderto measure the accurate conductivity of the fixed DNA. Recently, studieson fixation of DNA to solid bases have been reported by various methodssuch as ion connection, covalent bond, and protein bonding for the useof DNA as a potential electronic material.

[0151] Okahata, et al. prepared a polyion complex using DNA and cationlipids in order to prepare thin cast film membranes of DNA (Ijiro, K andOkahata, Y: A DNA-Lipid Complex Soluble in Organic Solvents. J. Chem.Soc., Chem. Commun. 1992:1339, 1992). Phosphate and cation lipids formedquantum chemical ion pairs. As a result, an alkyl base covered the DNAforming the shape of a brush to wash a test tube and became hydrophobicand settles instantly. Nishi et al. prepared the gel film with thethickness of 2-3 μm×2-3 mm by adding bivalent metallic ions such as Ca²⁺or Mg²⁺ to a water solution of alginic acid, a polysaccharide having aresidue of alginic acid (Iwata, K; Nishi, N; Miura, Y; Nishimura, S;Tokura, S: Polymer Preprints, 42:599, 1993). DNA structure wasmaintained in the film from adsorption test of intercalator color in thestudy. The molecular orientation of DNA in the film prepared by fixationmethods was random and was very difficult to control the molecularorientation and thickness of the membrane. G. Decher et al. reported onthe methods for preparing the thin membrane of DNA which had a thicknessof one molecule (Lvov, Y; Decher, G; Sukhorukov, G: Assembly of ThinFilms by Means of Successive Deposition of Alternate Layers of DNA andPoly(Allylamine). Macromolecules 26:5396-5399, 1993). High molecularweight DNA isolated from sturgeon sperm formed layers 33 Å thick byx-ray diffraction indicating the DNA spread two-dimensionally with thelong axis parallel to the substrate surface. In conventional studies,fixation was performed using the ion connection of anion phosphates atmultiple points. On the other hand, Maeda et al. reported the fixationmethods fixed the special edge of DNA on a gold terminal by chemicallytreating the edge of DNA with a thiol base (Maeda, M; Nakano, K; Uchida,S; Takagi, M: Mg²⁺-Selective Electrode Comprising Double-Helical DNA asReceptive Entity. Chem. Lett. 1994:1805-1808, 1994). Organic thiolcompounds bind strongly to gold. Maeda et al. considered that theorientation of DNA was vertical towards the terminal from themeasurement of the amount of fixed DNA. Ijiro et al. reported aproduction of a semi-molecular membrane using DNA, a cation intercalatorlipid (C₁₈-acridine orange), and Langmuir-Blodgett techniques of castinga thin film. Orientation of the DNA strings was attempted by applyingcompression and measuring conductivities in different directions (Ijiro,K; Shimomura, M; Tanaka, M; Nakamura, H; Hasebe, K: Thin Solid Films (inpress). Ijiro, K and Shimomura, M: Double-Stranded DNA for MolecularElectronic Devices. Kotai Butsuri 30(12):1042-1048, 1995. Birdi, K S:Lipid and Biopolymer Monolayers at Liquid Interfaces, Plenum Press; NewYork, London; 1989). As evidenced by this review of various methods forfixation of DNA on surfaces, there is some difficulty in orienting DNAfilms for use as routine commercial electronic materials providing highdensity molecular wire interconnects on common semiconductor or standardelectrical component substrates.

[0152] In a preferred embodiment of this invention, DNA or nucleic acidis used as the conducting polymer precursor to be electrochemicallydeposited and uniaxially oriented into a highly conductive liquidcrystalline form on the semiconductor substrate surface. Single-strandedDNA is not electrically conductive as a molecular wire. It is a randomcoil with little order. However, double-stranded A-, B-, or Z-DNA areexamples of flat heteroaromatic purine and pyrimidine i-stacked basepairs (i.e., heteroaromatic π-stacking of flat base pairs, one on top ofthe next in a rising helix) that makes double-stranded DNA conductive.Other examples of suitable DNA structures that may be deposited asuniaxially oriented liquid crystalline DNA quantum wires include, butare in no way limited to clockwise double-stranded twining structures,otherwise called A-, B-, C-, D-, E-, and T-types. DNA also has acounterclockwise double-stranded twining structure, called Z-type. Inaddition, there is looped DNA which consists of thousands of pairs ofbases called plasmid DNA which exists in prokaryotic organisms. There isalso a twisted looped DNA structure which comprises several loops and asuper helical structure. There even exists a twisted loop, cross shapedDNA (Ijiro, K and Shimomura, M: Double-Stranded DNA for MolecularElectronic Devices. Kotai Butsuri 30(12):1042-1048, 1995). And DNAexists in triple helix type structures as well (Povsic, T J; Dervan, PB: Triple Helix Formation By Oligonucleotides On DNA Extended To ThePhysiological pH Range. J. Am. Chem. Soc. 111(8):3059-3061, 1989).

[0153] Preferably, a liquid crystal B-DNA type double-stranded structureis deposited, electrically attached, and uniaxially oriented in parallelextended conformation orthogonal to the surface of a semiconductor inspecific chemically or electrochemically activated regions (as shown inFIG. 2). A and T; G and C complementary pairs of bases form an uprightduplex helical structure with a diameter of approximately 20 Å,comprising two high molecular chains. The pitch of the duplex helicalstructure is approximately 34 Å and 10 of the pairs of bases line upvertically towards the extended line of DNA. The upper and lower pairsof bases create an angle of 36° while the distance between each pair ofbases is 3.4 Å. This produces a strong mutual relationship between eachstuck pair of bases inside the duplex helical structure of DNA. Forexample, an extreme reduction of absorbance (light color effect) willoccur because of π-π*conversion. In other words, the internalcharacteristics of DNA can be considered as a suspected one-dimensionalcrystalline structure of stuck pairs of bases (Ijiro, K and Shimomura,M: Double-Stranded DNA for Molecular Electronic Devices. Kotai Butsuri30(12): 1042-1048, 1995).

[0154] Particularly high packing efficiencies are achieved in theicosahedral double-stranded DNA bacteriophages, where the DNA duplexesare close packed at a center-to-center spacing of about 26 Å. Thisconstraint has been incorporated into several recent models in all ofwhich the rods of duplex DNA are configured in more-or-less parallelbundles (Booy, F P; Newcomb, W W; Trus, B L; Brown, J C; Baker, T S;Steven, A C: Liquid-Crystalline, Phage-Like Packing Of Encapsidated DNAIn Herpes Simplex Virus. Cell 64:1007-1015, 1991). Moreover, the average26 Å interduplex spacing closely resembles that observed for liquidcrystals of DNA in vitro by cryoelectron microscopy or x-ray diffraction(Booy, F P; Newcomb, W W; Trus, B L; Brown, J C; Baker, T S; Steven, AC: Liquid-Crystalline, Phage-Like Packing Of Encapsidated DNA In HerpesSimplex Virus. Cell 64:1007-1015, 1991). In a preferred embodiment ofthis invention, uniaxially oriented liquid crystalline B-DNA conductivewires are electrochemically deposited at specific light activatedregions on the surface of a p-n junction solar cell by EMOLE fabricationmethods as described above.

[0155] V. Molecular Recognition Surfaces

[0156] A molecular recognition surface preferably is made up of atwo-dimensional crystal array of one or more molecular recognitionsite(s) that recognize a particular ligand (i.e., analyte) typically,though not necessarily, in a liquid. In addition to its ability to bindspecific ligands, a molecular recognition site may also be a catalyticsite, redox site, electron transfer site, energy transfer site, magnetictransfer site, and as a consequence of ligand binding may induceconformational change, and quantum-confined electron/hole tunneling andpercolation.

[0157] The molecular headgroups employed in this invention include, forexample, proteins (which bind ligands), catalytic antibodies,porphyrins, lectins, enzymes (including any enzyme categorized in the ECNomenclature—e.g., class 1: oxidoreductases, class 2: transferases,class 3: hydrolases, class 4: lyases, class 5: isomerases, and class 6:ligases), immunological antibodies, antigens, receptors, viruses, cells,cavitands, zeolites (which bind redox metals), supramolecularassemblies, electro-optical materials (e.g., second- and third-ordernonlinear optical materials), photoconductive and photoelectricmaterials (in which an applied electromagnetic field produces freeelectrons), giant magnetoresistive materials (in which an appliedmagnetic field changes resistivity of the material), metal chelates,magnetic materials (in which magnetic ordering is changed by thepresence of other magnetic materials), inorganic scintillators (whichconvert high energy radiation to lower energy light photons), inorganiccrystal oscillators (which act as a quantum frequency transmitter andreceiver), piezoelectric materials (in which mechanical force produceselectron flow), light-harvesting polymer systems (in which lightproduces electron flow and chemical energy storage), laser switch dyes(which absorb light at one wavelength and emit a monochromatic light ata longer wavelength), barrier tunnel switches (e.g., molecular electronswitches), etc.

[0158] Examples of ligands that can be used with this invention include,but are not restricted to, agonists and antagonists for cell membranereceptors, toxins and venoms, viral epitopes, antigenic determinants,monoclonal and polyclonal antibodies, hormones, hormone receptors,steroids, peptides, enzymes, substrates, cofactors, drugs, lectins,sugars, oligonucleotides, oligosaccharides, proteins, transition metals,chelates, cavitands, pollutants, chemical and biological warfare agents,poisons, dyes, gases, intercalators, alcohols, alkaloids, fats, lipids,cholesterol, blood type, cell surfaces, metabolites, etc.

[0159] Molecular recognition sites that mediate a biological or chemicalfunction either directly or indirectly on binding with a particularligand(s) are of most interest. Suitable molecular recognition sitesinclude relatively small, single molecules, such as cofactors, whichshow specific binding properties. Typically, molecular recognition siteswill range from 1 dalton to greater in size. Other examples of molecularrecognition sites include, but are not restricted to, the common classof receptors associated with the surface membrane of cells and include,for instance, the immunologically important receptors of B-cells,T-cells, macrophages and the like. Other examples of molecularrecognition sites that can be investigated by this invention include butare not restricted to hormone receptors, hormones, drugs, cellularreceptors, membrane transport proteins, electron transport proteins,steroids, peptides, enzymes, substrates, cofactors, vitamins, lectins,sugars, oligonucleotides, intercalators, oligosaccharides, viralepitopes, antigenic determinants, glycoproteins, glycolypoproteins,immunoglobins, restriction enzymes, catalytic antibodies, transitionmetals, chelates, cryptands, cavitands, supramolecular structures, etc.

[0160] A. Oxidoreductases (Redox Enzymes)

[0161] Examples of molecular recognition sites that bind specificligands, catalyze a redox reaction, and are electrically conductingbiopolymers, are a broad class of enzymes called the oxidoreductases. Tothis class belong all enzymes catalyzing oxido-reductions. The substrateoxidized is regarded as hydrogen or electron donor. The classificationis based on ‘donor:acceptor oxidoreductase’. The recommended name is‘dehydrogenase’, wherever this is possible; as an alternative, ‘acceptorreductase’ can be used. ‘Oxidase’ is only used in cases where O₂ is anacceptor. Classification is difficult in some cases because of the lackof specificity towards the acceptor. The EC number 1.x.x.x as it appearsin Enzyme Nomenclature (1978) is assigned to the class calledoxidoreductases (Enzyme Nomenclature. Academic Press; New York; 1978).

[0162] Oxidoreductases or redox enzymes are molecules of 40,000 daltons(e.g., galactose oxidase) to 850,000 daltons (e.g., cholinedehydrogenase) with one or more redox centers. Their averagehydrodynamic diameters range from ˜55 to ˜150 Å. In the great majorityof enzymes, the redox centers are located sufficiently far from theoutermost surface (defined by protruding protein or glycoproteindomains) to be electrically inaccessible. Consequently, most enzymes donot exchange electrons with electrodes on which they are adsorbed, i.e.,their redox centers are neither electrooxidized at positive potentialsnor electroreduced at negative ones. Apparently, part of the protein orglycoprotein shell surrounding the redox centers is there to preventindiscriminate electron exchange between the different redoxmacromolecules of living systems. Another function of this shell is tostabilize the structure of the enzyme. Because neither function isessential for catalysis, redox enzymes do function when part of theshell is stripped or, when the shell is chemically altered so as to makeit electrically conductive.

[0163] Examples of oxidoreductase enzymes suitable for use with thisinvention include glucose oxidase, catalase, peroxidase, cholesteroloxidase, and alcohol dehydrogenase. Glucose oxidase (GOD) turns over atambient temperature at a rate of ˜10² s⁻¹, i.e., it produces about 200transferable electrons/s. Because it has a radius of ˜43 Å, there can beup to 1.7×10¹² enzyme molecules on the electrode surface. The currentdensity, when all redox centers are electrically well connected to theelectrode, may thus reach about 3.4×10¹⁴ electrons s⁻¹ cm⁻², or 53 uAcm⁻².

[0164] In a preferred embodiment, molecular recognition site(s) will becomposed of one or more of the following oxidoreductases (redoxenzymes): glucose oxidase (GOD) which binds specifically to D-glucose,cholesterol esterase/cholesterol oxidase (COD) which binds specificallyto cholesterol ester/cholesterol, catalase (CAT) which bindsspecifically to H₂O₂, or alcohol dehydrogenase (ADH) which bindsspecifically to ethanol. All of these redox enzymes oxidize theirrespective substrates, transferring two electrons to natural orartificial diffusible electron acceptor mediators. In the presentinvention, a uniaxially oriented liquid crystal conducting biopolymer inan extended straight conformation is stuck or “wired” into eachcatalytic site/redox center permitting direct electron transfer to takeplace. Electron transfer to natural diffusible electron acceptors suchas O₂ or other artificial diffusible redox mediators such as ferroceneor metal derivatives is therefore largely eliminated. Mechanism ofelectron transfer in the present invention is based on a solid-state“hard-wired” organization at the enzyme catalytic site/redox centerestablishing quantum-confined electron/hole tunneling and percolationthrough a uniaxially oriented liquid crystal conducting polymer orbiopolymer known as a molecular or quantum wire. Electron or holeinjection from a molecular recognition headgroup (i.e., oxidoreductase)through an attached superconducting quantum wire tail (i.e., DNA)interconnect to an underlying electronic substrate is the basis of amolecular transistor.

[0165] In a preferred embodiment of this invention, a plurality of suchmolecular recognition sites (i.e., enzymes) are electrochemicallydeposited onto the surface of a p-n junction solar cell by firstdepositing liquid crystalline highly oriented B-DNA “molecular wires” tothe p-type surface. Preferably, a liquid crystalline molecularrecognition surface structure is deposited, electrically attached, anduniaxially oriented at the surface of a liquid crystal B-DNAdouble-stranded structure which was deposited, electrically attached,and uniaxially oriented at the surface of p-type semiconductor inspecific chemically or electrochemically activated regions. Oriented DNAduplex polyelectrolytes likely are extended, straight quantum wires thatpenetrate deeply into enzyme crevices at one end and semiconductorsubstrate at the other end. This type of molecular-scale structurelikely facilitates direct, quantum mechanical electron transfer betweenenzyme headgroups and semiconductor substrate.

[0166] In a preferred embodiment, spatially addressable electrochemicalactivation at specific regions on the surface of a p-n junction solarcell is achieved by light masking or photolithographic techniques forthe purpose of electrodeposition at specified locations on the chip. Ina preferred embodiment of this invention, liquid crystalline highlyoriented molecular recognition surfaces are electrochemically depositedat specific light activated regions on the surface of a p-n junctionsolar cell by EMOLE methods as described above. Preferably, DNA wires onthe p-n junction solar cell are exposed to light at specific regions toform electrical contacts with liquid crystal oriented molecularrecognition sites by EMOLE methods. This is repeated at differentregions on the semiconductor surface to pattern complex digital organicintegrated circuits (IC) of “wired” molecular recognition sites. Thefabrication scheme described above constitutes preferable productionmethods of a molecular recognition chip (MRC).

[0167] B. Immunoglobulins

[0168] If we are looking for a more general method of incorporatingnon-biological molecules into molecularly organized materials, then theimmunoglobulins or antibody molecules offer many attractive advantages.Using currently available monoclonal antibody technology, it is nowpossible to generate a specific immunoglobulin molecule capable ofbinding to almost any compound of interest. In accordance with thepresent invention, one could engineer crystals of antibody complexes inwhich it was possible to control the arrangement and orientation of thecomplexed molecules at the molecular-scale. There has already been areport of successful application of Langmuir-Blodgett techniques toproduce two-dimensional crystals of antibody molecules which may be usedfor MED development.

[0169] Examples of molecular recognition sites that bind specificligands, catalyze a redox reactions, undergo conformational change andare electrically conducting biopolymers, are a broad class of proteinscalled immunoglobulins. Catalytic antibodies are man-madeimmunoglobulins that can be engineered to possess all of the abovechemical and physical properties and specificity for a particularligand. In a preferred embodiment, immunoglobulins or catalyticantibodies may be deposited as molecular recognition headgroups onto DNAquantum wires using EMOLE crystallization processing techniques asdescribed above to fabricate molecular recognition (MR) devices on amacro-solid substrate.

[0170] VI. Conduction Mechanisms through Polymers on Solid Substrates

[0171] A. Energy Bands in Uniaxially Oriented Liquid Crystal ConductingBiopolymers (Proteins and DNA) and Semiconductor Substrates

[0172] Since Szent-Gyorgyi's report that biopolymers can work likesemiconductors, many researchers have pursued research on electronmovement through proteins. The potential for long-range electronmovement within a protein coupled with double helix DNA wastheoretically calculated from the point of quantum chemistry. Becauseionic impurities are present in DNA, the methods used to prepare solidpellets varied depending on the experiments and thus, reportedconductivities have varied between 10⁻⁴ and 10⁻¹⁰ mho·m⁻¹. A quantummechanical-based model also offers a possible explanation for theanomalously rapid long-range (i.e., ˜40 Å) photoelectron transferrecently observed by Barton and Turro et al. for donor and acceptorspecies intercalated into a DNA double helix.

[0173] There is no possibility of intrinsic conductivity in periodic andaperiodic polypeptide chains due to their large fundamental energy gap.This conclusion may appear, at first glance, to be a stumbling block tothe electronic conduction in proteins. It should however be noted thatmany other materials, the glasses, oxides and amorphous semiconductors,also have energy gaps sufficiently large to make them poor conductorsbut this has not prevented consideration of them in electronic terms andthe establishment of a considerable body of experimental and theoreticalevidence for long range electron transfer in them.

[0174] Since the bands in the density of states (DOS) curves ofaperiodic chains are very broad with a few small gaps, there is apossibility of extrinsic conduction on doping with electron acceptors(p-doping) or with electron donors (n-doping) in these chains. To decideabout the nature of extrinsic conduction (whether Bloch-type conductionor charge transport through hopping) one needs to investigate thelocalization properties of the wave functions belonging to the energylevels in the upper part of the valence band region or the lower part ofthe conduction band region (these are the regions of interest if acharge transfer is to take place in vivo due to the interaction ofproteins with electron acceptors or donors or with DNA). The possibilityof this type of charge transfer has also been suggested bySzent-Gyorgyi.

[0175] Quantum mechanical models proposed to estimate energy bands andelectronic conduction in proteins and DNA can be influenced by a numbersof external factors which tend to reduce or eliminate the bandgap andbroaden the width of estimated valence and conduction bands. This yieldsbiopolymers with metallic-like conduction properties. Such externalfactors include impurities, dopants, applied electric fields, appliedmagnetic fields, illumination (hν), hydration with H₂O, solvent,pressure, conformational changes, orientation, pH, electrolytes, localsurface charges, and injection of electron or holes directly into theconduction or valence bands of the biopolymer. Injection of electronsinto protein conduction bands can come from COO⁻ groups on protein sidechains or at the carboxyl terminus, and from H₂O. Selective applicationof these external factors effects are used to engineer bandgap structureof proteins and DNA using EMOLE fabrication techniques to producedesired physical and chemical properties of superconducting, conducting,semiconducting, or insulative forms. EMOLE provides energy band matchingand molecular interconnects between proteins, DNA, and the semiconductorsubstrate which affords quantum mechanical electronic conduction.

[0176] Uniaxially oriented liquid crystalline forms of conductingbiopolymers (proteins and DNA) may be produced by EMOLE fabricationtechniques. Processing variables utilized by EMOLE to deposit orientedliquid crystal conducting biopolymers include external factorsinfluencing biopolymer energy band structures described above. EMOLE isa chip fabrication method used to engineer molecular structure, energyband structure, band matching, and quantum mechanical molecularinterconnects of conducting biopolymers (proteins and DNA) on thesurface of a semiconductor substrate.

[0177] In order for communication between uniaxially oriented liquidcrystal conducting biopolymers (proteins and DNA) and thepolycrystalline or monocrystalline macro-semiconductor substrate of theMR-device, common energy levels must exist between not only the protein(molecular headgroup) and DNA (quantum wire tail) components, butbetween the DNA and the semiconductor substrate. In MR-devices of thepresent invention, DNA duplex polyelectrolytes are extended, straightquantum wires that penetrate deeply into enzyme crevices at one end, andinto the macro-semiconductor substrate at the other end. This type ofmolecular-scale structure facilitates direct electron transfer from theenzyme prosthetic group and desired energy continua between enzyme, DNA,and semiconductor substrate. The nature of the energy continua issimilar to the ideas proposed by Szent-Gyorgyi in 1946, Pethig,Bathaski, Tanatar, and others regarding a common quantum mechanicalenergy band continuum, resonant tunneling, hopping, acoustic plasmon,etc. mechanisms which facilitate charge transfer of mobile chargecarriers (electrons or holes) through protein and DNA to the underlyingsemiconductor substrate.

[0178] B. Superconductivity

[0179] The possibility that superconductive phenomena may play abiological role is at present a controversial subject in severallaboratories. Unlike the situation for normal electronic conductors,electrons in a superconductor are not free to move independently of eachother but exist as coupled electron pairs constrained to be in the samequantum state. As a result of this pairing-up of electrons, electronscattering effects are minimized with the result that the flow ofelectron current can occur without the generation of heat and hence withno electrical resistance. Such an effect could obviously havefar-reaching consequences if it could be detected in biological systemsat physiological temperatures. In conventional superconductors theelectron pairing results from interactions between the electrons and thelattice phonons. In 1964, Little proposed that suitably constructedorganic polymeric systems would be capable of sustainingsuperconductivity as a result of an electron-pairing mechanism involvingelectron-exciton interactions (Little, W A: Possibility of Synthesizingan Organic Superconductor. Phys. Rev. 134(6A):A1416-A1424, 1964.).Little estimated that such a polymer, consisting of a conductingconjugated hydrocarbon backbone and side chains in the form of highlypolarizable dye molecules, would be superconducting up to temperaturesof the order 2200° K. Such high temperatures would obviously not berealistic for organic systems for reasons of thermal stability, but thisestimate of the critical temperature does serve to indicate that theconcept of the existence of superconducting biopolymers at physiologicaltemperatures lies well within the limit of the applicability of Little'stheory. The existence of superconductivity in aromatic compounds wasfirst speculated upon by London (London, F J: J. Phys. Radium 8:397,1937); and Ladik et al. (Ladik, J; Biczo, G; Redly, J: Possibility ofSuperconductive-Type Enhanced Conductivity in DNA at Room Temperature.Phys. Rev. 188(2):710-715, 1969) have provided a theoretical basis forthe superconductive behavior of DNA.

[0180] Experimental evidence for high temperature superconduction inbiological molecules has been reported in several laboratories.Superconductivity was deduced to occur in small domains included in theinsulating bulk of bile cholate test samples and so as to distinguishthe effects from that normally found for the elemental superconductors,the cholates were designated a fractional or Type III superconductor.When small amounts of water are introduced into such materials thehydrophobic groups will tend to cluster together, and on subsequent slowdesiccation small micelles will be formed. Such micelles are consideredby Halpern and Wolf to form superconducting domains.

[0181] Following the suggestion that enzymes and other biologicalmaterials possess a metastable state with high dipole moment, Ahmed etal. investigated the dielectric and magnetic susceptibility propertiesof the dilute solutions of lysozyme (Ahmed, N A G; Calderwood, J H;Frohlich, H; Smith, C W: Evidence For Collective Magnetic Effects In AnEnzyme: Likelihood Of Room Temperature Superconductive Regions. Phys.Lett. 53A(2):129-130, 1975). It was found that magnetic fields of theorder of 0.6 tesla could produce very large changes (˜30%) in therelative permittivity of the solutions. This was suggestive ofsuperconductive behavior. It was suggested that in each lysozymemolecule there existed a small superconductive region with lineardimensions smaller than the London penetration depth, and that thecollective, superconductor-like, phenomena resulted from the formationof clusters of these small regions. This is similar to the cluster modelproposed for bile cholates. It was also suggested that not only thelysozyme molecules, but also water and ions may have played a role inthe establishment of the superconducting regions.

[0182] Other indirect evidence to suggest a biological role forsuperconductivity has been suggested by Cope (Cope, F W: Physiol. Chem.Phys. 3:403, 1971. Cope, F W: Physiol. Chem. Phys. 5:173, 1973) thathigh temperature superconduction may be expected in a sandwichconsisting of a thin conductive film or filament adjacent to adielectric layer. Cope considers that such superconducting sandwichesmay be ubiquitous in biological systems in the form of thin layers ofprotein and unsaturated lipids and hydrocarbon ring structures(conducting layer) adjacent to layers of water (polarizable dielectriclayer). Examples of such biological processes are impulse conductionvelocity in frog sciatic nerves and junctional electrical resistance ofcrayfish nerve. Such an effect can be well described in terms of a modelwhere the rate-limited biological process involves a superconductingtunneling current of single electrons and/or electron pairs (theJosephson current). It was suggested that as there was an apparentassociation of superconduction with growth, then the superconductivemicro-regions may have been individual purine and pyrimidine rings ofDNA and RNA with electron tunneling between rings along the length ofthe polymer chain. It was further suggested that superconductiveJosephson junctions in living systems may provide a physical mechanismwith more than enough sensitivity to explain how many biologicalorganisms are able to respond to weak magnetic fields.

[0183] Two-component plasmas (or more generally multi-component plasmas)as in an electron-hole liquid can support, other than the usual plasmonmode, a new collective mode called the “acoustic-plasmon mode”. Quantummechanical treatment of acoustic plasmons in one dimensional systemssuch as a long DNA molecule have attracted attention. Tanatar (Tanatar,B: Collective Modes in a Quasi-One Dimensional, Two-Component ElectronLiquid. Solid State Communications 92(8):699-702, 1994) stated that amotivation to study the acoustic plasmons in quasi-one-dimensionalelectron-hole systems comes from the fact that they may provide apairing mechanism like the BCS theory which leads to a superconductingtransition (Bardeen, J; Cooper, L N; Schrieffer, J R: Microscopic Theoryof Superconductivity. Phys. Rev. 106:162-164, 1957). Such an acousticplasmon mediated superconductivity has been proposed and elaborated fortwo-dimensional electron-hole liquids. Possibility of superconductivitydue to ordinary plasmons in quantum wires were also considered.Experiments to observe the acoustic plasmons in quasi-one-dimensionalstructures such as DNA, and their possible pairing mechanism leading tosuperconductivity would be most interesting.

[0184] In a preferred embodiment, uniaxially oriented liquid crystalconducting biopolymers (protein and DNA) deposited by controlled EMOLEfabrication techniques are used to produce a functional device. Suchdevices are thought to function via one or more superconductingmechanism(s) described above. For example, a GOD-DNA device generates anelectron pair for each D-(+)-glucose molecule oxidized by the GODprotein enzyme headgroup. Electron pair movement from protein FAD/FADH₂prosthetic group (redox center) through DNA quantum wire to underlyingsemiconductor substrate occurs via superconducting mechanism(s)described above. Many gated devices inject an electron pair, viasuperconducting mechanism, into p-type silicon of a p-n homojunctionsolar cell; combining with photogenerated majority carriers (holes), tolower the baseline photocurrent (I_(sc)). Decrease in photocurrent isdirectly proportional to D-(+)-glucose concentration. Change inphotocurrent occurs very rapidly and is accompanied by a near stepchange (see FIGS. 5 and 6 described below), resulting from differentialdevice injection of mobile charge carriers (electrons or holes) intop-type or n-type semiconductor substrate surfaces.

[0185] VII. Applications

[0186] The sensors of this invention may be employed for a myriad ofapplications. For example, sensor based home health monitors will besimple-to-use, non-invasive and relatively inexpensive for use inmonitoring health conditions at home. Many physical functions—liverfunctions, ovulation, pregnancy, yeast infections, viral infections,bacterial infections, levels of cholesterol, triglycerides, sugar,hormones, drugs, water, salt, pH, sodium, and potassium—may be monitoredas easily as weight is now tracked by bathroom scales. The graying ofour population and the increasing costs of medical care will make theseproducts extremely popular.

[0187] In another preferred embodiment of this invention, a sensor, in aportable pen-based device, may be used to monitor compounds found in thehuman breath. The normal human breath contains hundreds of volatileorganic compounds that are reflective of the metabolic state of theperson. These volatile organic compounds have been quantitated by gaschromatographic (GC) and mass spectrometry (MS) methods in numerousstudies. Preferably, a sensor of this invention is exposed to exhaledbreath. In a preferred embodiment, the molecular recognition surface ofthe sensor will be alcohol dehydrogenase (ADH) which specifically bindsto ethanol; reduced mercaptoethanol, glutathione, or dithiothreitolwhich specifically binds to sulfur containing compounds; or a variety ofother molecular recognition sites to detect breath compounds readilyrecognized by those of skill in the art. The ADH-sensor will providepolice and highway patrol officers with a portable pen-basedbreathalyzer to validate drunk driving violations in the field. TheThio-sensor will provide individuals with a portable pen-basedbreathalyzer for discrete detection of halitosis (i.e., bad breath).

[0188] In another preferred embodiment, a device-based molecularrecognition chip (MRC) may be embedded in a magnetoosmotic (MOP) orelectroosmotic patch (EOP) which may be applied to the skin forreal-time non-invasive quantitation of analytes found below the skin(i.e., analytes in blood and deep anatomic structures). This is anon-invasive approach to analyte quantitation alternate to exposure ofthe powered chip to invasively drawn blood or other fluids describedabove. The MRC-MOP or MRC-EOP is suitable for non-invasive detection ofsmall charged, uncharged, and zwitter ionic molecules and salts (i.e.,analytes) less than 30,000 daltons found on the other side of complexsynthetic or biological barriers such as skin, adipose tissue, vascularwalls (i.e., venous and arterial vessel walls), isoparenteral walls,extravascular walls, extracellular walls, cerebral vascular walls, bloodbrain barrier (BBB), and a variety of other man-made and naturalmembranes. The MOP, applies a combination of localized magnetic fieldgradients and hypertonic junctions to surfaces such as skin that itcontacts. The EOP, applies a combination of localized electric fieldgradients and hypertonic junctions to surfaces such as skin that itcontacts. This permits the MOP or EOP to draw analytes throughsemi-permeable membranes and skin for detection by the embedded MRC asdescribed above. Preferably, the MRC-MOP or MRC-EOP may be equipped witha number of molecular recognition sites to perform a complete blood gas,blood electrolyte, hematocrit, blood sugar, and blood metaboliteanalysis non-invasively (i.e., without drawing blood).

[0189] In a preferred embodiment, applied a.c or d.c. electric ormagnetic fields are utilized to change the orientational and positionalorder of liquid crystal biological structures such as cellularmembranes, cellular pores, blood vessels, skin, sweat glands, etc. topermit leakage of contained body analytes. A hypertonic junction willpull out, by means of a low chemical potential well, and concentrateleaky analytes. The hypertonic junction is composed of a suitablepolyelectrolyte gel or solid polymer electrolyte (Gray, F M: SolidPolymer Electrolytes. Fundamentals and Technological Applications. VCHPublishers, Inc.; New York, Weinheim, Cambridge; 1991. Hara, M (ed.):Polyelectrolytes. Science and Technology. Marcel Dekker, Inc.; New York,Basel, Hong Kong; 1993) containing an embedded device-based molecularrecognition chip (MRC) for detection of specific analyte(s).

[0190] VIII. Screening and Assays

[0191] A semiconductor surface prepared according to the methodsdescribed above can be used to screen for ligands (i.e., analytes)having high affinity for immobilized molecular recognition sites. Asolution containing an unmarked (not labeled) ligand is introduced tothe surface. Generally, little or no incubation time is required becauseof the immediate response of the molecular recognition chip (MRC) on theorder of milliseconds.

[0192] In a preferred embodiment, a semiconductor substrate prepared asdiscussed above is exposed to light while connected to a digitalmultimeter (DMM) which measures the short circuit volt/amp output (i.e.,V_(sc), I_(sc)) of the p-n junction solar cell substrate (as shown inFIG. 3). The powered chip can now be exposed to a solution containing anunmarked ligand. The unmarked ligand binds with high affinity to animmobilized molecular recognition site previously localized on the chipsurface. A square wave signal is generated by the powered chip in lessthan a few milliseconds in response to the binding event of the ligand(i.e., digital output from the chip). In a preferred embodiment,D-(+)-glucose is applied to the surface of a glucose oxidase-molecularrecognition chip (GOD-Chip). The light powered GOD-Chip produces asquare wave response in volt/amp output proportional to the appliedD-(+)-glucose concentration (see FIGS. 5 and 6 described below). Thisreflects a change in the short circuit output of the p-n junction solarcell substrate due to electron tunneling from the molecular recognitionsurface (i.e., GOD) through the highly conductive polymer monolayer(i.e., liquid-crystal oriented B-DNA) to the p-type surface of the p-njunction solar cell. Carrier injection of electrons by a device into thep-type layer of a powered solar cell substrate interrupts the baselineshort circuit photovoltage and photocurrent (i.e., V_(sc), I_(sc)) ofthis simple p-n junction, rectifying diode device. The amount ofelectrons injected into the p-type surface of the powered chip isproportional to the amount of D-(+)-glucose binding to the GOD, which isreflected in a proportional digital square wave output by the GOD-Chip(FIG. 5 and FIG. 6). Device injection of electrons directly into p-typesilicon eliminates or lowers the photocurrent by combination withphotogenerated carrier holes before than can recombine, via shortcircuit wire, with photogenerated carrier electrons from the n-typelayer. Concomitantly with lowered photocurrent resulting from carrierhole removal from the p-type layer, the continued build-up ofphotogenerated carrier electrons in the n-type layer increases measuredphotovoltage of the circuit (FIG. 6).

[0193] In this embodiment, a simple digital multimeter (volt/amp) isemployed to measure the digital output of the GOD-Chip. Therefore,single and multiple IC arrays described above may be configured inpen-based digital meters, hand-held digital meters, clinical lab-basedinstruments, digital wireless implantable medical devices, andindustrial-based digital devices which measure real-time molecularbinding events and constants of analytes. A simple calibration curve foreach chip can be used to determine the concentration of unknown samples.Calibrated chips are not affected by altitude, humidity, O₂ partialpressure, diffusional electron acceptor mediators, or application of thesample. These problems of the prior art, have been overcome in thepresent invention because electron transfer rates of the molecular wireinterconnects are orders of magnitude greater than enzymatic reactionrates, and electron transfer rates of diffusional redox mediators suchas O₂ and other small molecule inorganic, organometallic, and organiccompounds used in amperometric detection methods. A forward electrontransfer rate constant (k_(f)>10⁷ s⁻¹ Å⁻²) may be very high because ofthe quantum-wire nature (i.e., defined electronic energy levels) of theconductive polymer interconnects. Connecting polymers may also bereversibly switched between conductive and insulative states byoxidation or reduction.

IX. EXAMPLES

[0194] The following examples of preferred embodiments of the presentinvention are presented by way of illustration only and do not suggestthat the above-described methods and compositions are in any way limitedby the specific examples set forth below.

Example A Preparation of Polycrystalline Silicon p-n Junction Solar Cell

[0195] A commercial polycrystalline silicon p-n junction solar cell chip(a 0.1799 g and 1.75 cm²) from Edmund Scientific, Barrington, N.J.08007-1380 (Stock Nos. 35,220 and 35,221) was exposed to 980 Lux lightintensity from a F15T8/CW Westinghouse bulb at 25° C. With the dark blueemitter surface facing the light source, the short circuit DC output ofthe dry solar cell substrate was measured by a digital multimeter (DMM)(Extech Instruments; Model No. 383273). Measured DC output was 121 mVand 98 uA. The solar cell chip was then washed with analytical reagentelectronic-grade solvents: i) acetone; ii) methanol; iii) 18 Mohm H₂O;and iv) methanol. It was allowed to dry in a dust-free environment. Thisprepared the semiconductor p-n type surfaces for electroplating.

Example B Electrodeposition of DNA Onto a Polycrystalline Silicon p-nJunction Solar Cell

[0196] A DNA electroplating solution was prepared using 18 Mohm sterilewater as the solvent. 0.1062 g of DNA (degraded free acid from Herringsperm) was added to 100 ml of water. The pH of the resulting solutionwas ˜2.00. The pH was adjusted to ˜7.00 with NaOH and HCl. No buffer wasadded to the electroplating solution. The final salt/electrolyteconcentration was <150 mM. 1.00 mL of methanol was added to the DNAelectroplating solution and mixed thoroughly. The dry solar cell chipfrom Example A produced a short circuit output of 152 mV and 136 uA whenexposed to 1700 Lux light intensity generated from two F15T8/CWWestinghouse bulbs at 25° C. The dry solar cell chip from Example A wassubmerged in the DNA electroplating bath at 25° C. with dark blueemitter surface exposed to 1700 Lux light intensity generated from twoF15T8/CW Westinghouse bulbs. After ˜5.50 hours, the solar cell chip wasremoved from the DNA bath and placed on a paper towel to dry. The darkblue emitter surface was exposed to the 1700 Lux light source during thedrying process which took ˜12.00 hours at 25° C. in air. DNAelectroplated on the back or silvery side of the solar cell chip (i.e.,p-type silicon) as evidenced by a white coating visible to the eye. Onthe dark blue emitter surface (i.e., n-type silicon) no significantcoating was observed. The pH of the DNA electroplating bath remained7.00 after the electroplating process was complete. The electrochemicalpotential at the plating surface was about 150 mV and the currentdensity was about 77 μA cm⁻².

Example C Electrodeposition of Glucose Oxidase (GOD) Onto a DNA-CoatedPolycrystalline Silicon p-n Junction Solar Cell

[0197] A glucose oxidase (GOD) electroplating solution was preparedusing 18 Mohm sterile water as the solvent. 0.0092 g of glucose oxidase(EC 1.1.3.4; ˜1,000 units) was added to 100 ml of water. The pH of theresulting solution was ˜6.00. No buffer or further adjustment of pH wasnecessary. 1.00 ml of methanol was added to the GOD electroplatingsolution and mixed thoroughly. Next, the DNA-coated polycrystallinesilicon p-n junction solar cell from Example B was submerged in the GODelectroplating bath with dark blue emitter surface exposed to 1700 Luxlight intensity from two F15T8/CW Westinghouse bulbs at 25° C. for ˜8.10hours. The solar cell chip was removed from the GOD bath and placed on apaper towel to dry. The dark blue emitter surface was exposed to the1700 Lux light source during the drying process which took ˜12.00 hoursat 25° C. in air. GOD electroplated on the back or silvery side of thesolar cell chip (i.e., p-type silicon) as evidenced by a yellow-orangeprecipitate visible to the eye. The yellow-orange GOD precipitate was inthe same area of the chip overlapping the white DNA precipitate fromexample B. On the dark blue emitter surface (i.e., n-type silicon) nosignificant coating was observed. The pH of the GOD electroplating bathremained ˜6.00 after the electroplating process was complete. TheGOD-DNA-Chip was removed from the light and put under parafilm toprotect and store until use. The electrochemical potential at theplating surface was about 150 mV and the current density was about 77 μAcm⁻².

Example D Detection of D-(+)-glucose on a GOD-DNA-Chip

[0198] Coating/electroplating of the solar cell chip from example A didnot change the electronic output characteristics of the device prior totesting with the D-(+)-glucose ligand.

[0199] The dry GOD-DNA-Chip from example C was placed with the silverGOD-DNA coated surface (i.e., p-type silicon) facing a F15T8/CWWestinghouse bulb. A red (positive) test lead of a digital multimeter(DMM) Extech Instruments; Model No. 383273) was connected to the darkblue emitter (i.e., n-type) surface and the black (negative) test leadwas connected to the p-type GOD-DNA coated surface facing the light(FIG. 3). The intensity of light was adjusted to produce a baselineshort circuit current of approximately −60 uA (FIG. 5). After severalminutes, a drop (˜0.100 mL) of a sterile D-(+)-glucose standard (63mg/dL) was placed on the powered GOD-DNA-Chip resulting in a largesquare wave amplitude change of approximately +51 uA reaching a newbaseline of approximately −8 uA (FIG. 5). This is consistent withapproximately 2.00×10¹⁷ glucose molecules being applied to the chip in a1 cm² area generating the maximum current expected from a monolayer ofwell connected GOD. Glucose oxidase (GOD) turns over at ambienttemperature at a rate of ˜10² s⁻¹, i.e., it produces about 200transferable electrons/s. Because it has a radius of ˜43 Å, there can beup to 1.7×10¹² enzyme molecules on the electrode surface. The currentdensity, when all redox centers are electrically well connected to theelectrode, may thus reach about 3.4×10¹⁴ electrons s⁻¹ cm⁻², or 53 uAcm⁻² (Heller, A: Electrical Wiring of Redox Enzymes. Acc. Chem. Res.23(5):128-134, 1990).

[0200] Another test of GOD-DNA-Chip performance at differentD-(+)-glucose concentration levels is demonstrated in FIG. 6. “Level 1”and “level 2” are sterile D-(+)-glucose standards (˜63 and 20 mg/dLrespectively). A drop of “level 1” D-(+)-glucose standard produces thefirst square wave; followed by washing with H₂O and application of thelower “level 2” D-(+)-glucose concentration. Square wave amplituderesponses are directly proportional to the D-glucose concentrationsapplied to the chip. Washing the GOD-DNA-Chip of ligand D-(+)-glucosewith H₂O returns the chip to its baseline voltage/current (FIG. 5 andFIG. 6).

Example E Electrodeposition of Glucose Dehydrogenase (GDH) Onto aDNA-Coated Polycrystalline Silicon p-n Junction Solar Cell

[0201] A DNA-coated polycrystalline silicon solar cell was prepared in amanner similar to that explained above in Examples A and B. Thedifferences were as follows:

[0202] 1. A commercial polycrystalline silicon solar cell chip 0.0280 gand 0.4059 cm² was used as the semiconductor substrate.

[0203] 2. The dry solar cell chip from 1 (above) produced a shortcircuit output of 43.55 mV and 36.35 microAmperes when exposed to 1700Lux light intensity generated from two F15T8/CW Westinghouse fluorescentbulbs at 25 degrees Centigrade.

[0204] 3. The dry solar cell chip was submerged into 300 microLiters ofthe DNA/EMOLE™ electroplating bath at 25 degrees Centigrade with thedark blue emitter surface of the chip exposed to 1700 Lux lightintensity generated by two F15T/CW Westinghouse fluorescent bulbs.

[0205] 4. After 38.75 hours, the solar cell chip was removed from theDNA/EMOLE™ electroplating bath.

[0206] 5. The DNA-chip was dried under 1700 Lux light intensitygenerated by two F15T8/CW Westinghouse fluorescent bulbs and an umbrellaof N₂ gas for 2.50 hours at 25 degrees Centigrade.

[0207] 6. The electrochemical potential at the plating surface of thesilicon semiconductor substrate was about 44 mV and the current densitywas about 89 microAmperes/cm⁻².

[0208] A glucose dehydrogenase (GDH) electroplating solution wasprepared using 18 Mohm sterile water as the solvent. 0.0046 g of glucosedehydrogenase (EC 1.1.1.119; 50 units) was added to 7.50 mL of water.The pH of the resulting solution was 6.728. No addition of buffer orfurther adjustment of pH was necessary. 75 microLiters of methanol wasadded to the GDH electroplating solution and mixed thoroughly. Next, thedry DNA-chip from above was submerged into 300 microLiters of theGDH/EMOLE™ electroplating bath with the dark blue emitter surface of thechip exposed to the 1700 Lux light intensity generated by two F15T8/CWWestinghouse fluorescent bulbs at 25 degrees Centigrade for 26.75 hours.The solar cell chip was removed from the GDH/EMOLE™ electroplating bath.The GDH-DNA-chip was dried under 1700 Lux light intensity generated bytwo F15T8/CW Westinghouse fluorescent bulbs and an umbrella of N₂ gasfor 5.00 hours at 25 degrees Centigrade. The GDH-DNA-chip was removedfrom the light and put in a desiccator box to protect and store untiluse. The electrochemical potential at the plating surface of the siliconsemiconductor substrate was about 44 mV and the current density wasabout 89 microAmperes/cm⁻².

Example F Detection of D-(+)-Glucose on a GDH-DNA-Chip

[0209] As in the GOD examples, EMOLE™ coating/electroplating of thesolar cell chip did not change the electronic output characteristics ofthe device prior to testing with the D-(+)-glucose ligand.

[0210] The dry GDH-DNA-chip from Example E was placed with the silverGDH-DNA coated surface (i.e., p-type silicon) facing a F15T8/CWWestinghouse fluorescent bulb. A black (negative) test lead of a digitalmultimeter (Hewlett-Packard Model 34970A) was connected to the dark blueemitter (i.e., n-type silicon) surface and the red (positive) test leadwas connected to the p-type GDH-DNA coated surface facing the light. Theintensity of the light was adjusted to produce a baseline short circuitcurrent of approximately +65 microAmperes (lower curve of FIG. 7) andbaseline potential of approximately +78 mV (upper curve of FIG. 7).After a few minutes, 5 microLiters of sterile D-(+)-glucose (60 mg/dL)in saline sodium phosphate buffer (1×SSP, pH 7.323) was dropped on theGDH-DNA-chip resulting in an immediate large square wave amplitudechange of approximately +6.5 microAmperes and +7.5 mV reaching newbaselines of approximately +71 microAmperes and +85 mV, respectively(FIG. 7).

[0211] The above examples employing GOD and GDH serve to illustrate theutility and wide applicability of the present invention. While both GOD(EC 1.1.3.4) and GDH (1.1.1.119) oxidize D-(+)-glucose toD-gluconolactone, they are very different enzymes.

[0212] GOD (EC 1.1.3.4) is widespread among fungi. GOD is a FADcontaining flavoprotein and glycoprotein with a molecular mass of160,000 daltons. GOD contains two moles of FAD cofactor per mole ofenzyme and 16% carbohydrate, the carbohydrate chains are not directlyinvolved in catalysis. The specificity of GOD is very high, thebeta-form of glucose is oxidized 157 times more rapidly than thealpha-form and of other substrates examined only 2-deoxy-D-glucose and6-deoxy-D-glucose were oxidized at rates greater than 10% of that ofD-glucose. O₂ is the natural electron acceptor of this enzyme producingH₂O₂.

[0213] The NAD(P)-dependent GDH (EC 1.1.1.119) occurs inphotoautotrophic prokaryotes such as strains of bacteria capable offorming on glucose in the dark. In addition to oxidizing D-(+)-glucose(alpha- and beta-forms), NAD(P)-dependent GDH also oxidizes D-mannose,2-deoxy-D-glucose, and 2-amino-2-deoxy-D-mannose. NAD(P)-dependent GDHis not a flavoprotein or glycoprotein and has unusual specificity; itdoes not oxidize aldopentoses and is completely inactive with NAD⁺ or O₂as electron acceptors. Instead it very specifically requires NAD(P)⁺ asits electron acceptor, producing NAD(P)H+H⁺. The molecular mass of theenzyme is approximately 230,000 daltons. Oxidation of D-mannose is arelatively unusual feature of aldose dehydrogenases obtained fromvarious biological sources.

[0214] If NAD(P)⁺ is not available in solution, GDH (EC 1.1.1.119) willnot oxidize D-(+)-glucose. NAD(P)⁺ was not added to the GDH-DNA-chiptest solution in Example F which nevertheless rapidly oxidized addedD-(+)-glucose indicating that the DNA molecular wire of this device hasreplaced diffusible NAD(P)⁺, not present in the test solution, as a“hard wired” conduit for direct electron transfer from the attachedcatalytic headgroup GDH enzyme to the silicon semiconductor substrate.

[0215] As mentioned above, GOD in its native state oxidizes D-glucosethrough its FAD/FADH₂ redox center. This involves two electrons and twohydrogen ions being transferred to the FAD prosthetic group which istightly bound to the enzyme. Normally, in the absence of a sensormediator, the GOD-FADH₂ complex is reoxidized by atmospheric oxygen(i.e., O₂) to GOD-FAD complex to complete the catalytic reaction cycle.GDH (EC 1.1.1.119), by contrast, is not a FAD containing flavoprotein.GDH (EC 1.1.1.119) oxidizes D-glucose through a different redox centerutilizing diffusible NAD(P)⁺ coenzyme in stoichiometric amounts which isbrought into play during the catalytic mechanism of oxidation andelectron transfer producing D-gluconolactone and NAD(P)H+H⁺. The reducedcoenzyme NAD(P)H is not recycled nor reoxidized by molecular oxygen(i.e., O₂) (as with GOD-FADH₂) so that enough expensive NAD(P)⁺ coenzymemust be added in the beginning to drive the biocatalytic oxidation ofglucose. Thus, glucose sensors relying on GDH (EC 1.1.1.119) are notsensitive to oxygen partial pressure, unlike GOD-based glucose sensors.Still further, the GOD and GDH amino acid sequences are completelydifferent. The GDH enzyme has a molecular mass of approximately 230,000daltons while the GOD enzyme has a molecular mass of approximately160,000 daltons. Thus, the above examples demonstrate that the inventioncan be applied to widely different molecular recognition headgroups.

[0216] Conclusion

[0217] Various references have been cited in this specification. Each ofthese references is incorporated herein by reference for all purposes.

[0218] The invention has been described primarily with reference to theuse of electrochemical deposition of liquid-crystal conductive polymersand molecular recognition surfaces, but it will be readily recognized bythose of skill in the art that other types of deposition, conductivewiring, and substrates can be used. Various forms of patternedelectrochemical and chemical deposition may be used. Many types of p-nhetero- or homojunction semiconductor substrates may be used. Thesubstrate may be powered by broad spectrum light, light emitting diodes(LED), lasers, solar radiation, uv radiation, vis radiation, infraredradiation, x-rays, gamma rays, radioactivity, thermally, or by anyexternal supplied nuclear or electromagnetic energy greater than thesubstrate bandgap to provide patterned areas of electrochemicaldeposition and to power the completed device.

[0219] It is understood that the above description is intended to beillustrative and not restrictive. Many embodiments will be apparent tothose of skill in the art upon reviewing the above description. Thescope of the invention should, therefore, be determined not withreference to the above description, but should instead be determinedwith reference to the appended claims, along with the full scope ofequivalents to which such claims are entitled.

What is claimed is:
 1. A sensor for sensing the presence of an analytecomponent, which sensor does not rely on redox mediators, the sensorcomprising: a substrate that reports reception of mobile charge carriersto an electronic circuit; and a plurality of active molecularrecognition groups wired to a surface of the substrate by molecularwires aligned in a non random orientation, wherein when the analytecomponent contacts the active molecular recognition surfaces, mobilecharge carriers are transferred directly to and through the molecularwires attached to said headgroup surfaces, without redox reaction in themolecular wires, to thereby allow the substrate to sense the presence ofthe analyte component.